Cardiac event detection in implantable medical devices

Abstract: A cardiac event detection system and method detects arrhythmic cardiac activity or other information from an electrogram signal of a heart. The system senses the electrogram signal through an electrogram lead (12), preliminarily processes (34) the signal, and converts it to a plurality of discrete digital signals, each of which represents the magnitude of the electrogram signal at a prescribed sample time. The discrete digital signals are applied to both a cardiac event detector (38) and a morphology detector (36). The morphology detector (36) detects selected changes in the morphology (shape) of the electrogram signal, wherein such changes automatically control (30) the sensitivity (gain and/or threshold) used to detect cardiac events. The occurrence of a prescribed amount of change in the detected morphology over time indicates the occurrence of a prescribed arrhythmic cardiac condition.

Description: The present invention relates to cardiac event detection. More particularly, the present invention relates to the detection and tracking of cardiac events and electrogram morphology. The major pumping chambers in the human heart are the left and right ventricles. The simultaneous physical contraction of the myocardial tissue in these chambers expels blood into the aorta and the pulmonary artery. Blood enters the ventricles from smaller antechambers called the left and right atria which contract about 100 milliseconds (ms) before the ventricles. This interval is known as the atrioventricular (AV) delay. The physical contractions of the muscle tissue result from the depolarization of such tissue, which depolarization is induced by a wave of spontaneous electrical excitation which begins in the right atrium, spreads to the left atrium and then enters the AV node which delays its passage to the ventricles via the so-called bundle of His. The frequency of the waves of excitation is normally regulated metabolically by the sinus node. The atrial rate is thus referred to as the sinus rate or sinus rhythm of the heart. Electrical signals corresponding to the depolarization of the myocardial muscle tissue appear in the patient's electrocardiogram. A brief low amplitude signal known as the P-wave accompanies atrial depolarization normally followed by a much larger amplitude signal, known as the QRS complex, with a predominant R-wave signifying ventricular depolarization. Repolarization prior to the next contraction is marked by a broad waveform in the electrocardiogram known as the T-wave. A typical implanted cardiac pacer (or pacemaker) operates by supplying missing stimulation pulses through an electrode on a pacing lead in contact with the atrial or ventricular muscle tissue. The electrical stimulus independently initiates depolarization of the myocardial (atrial or ventricular) tissue resulting in the desired contraction. The P-wave or R-wave can be sensed through the same lead, i.e., the pacing lead, and used as a timing signal to synchronize or inhibit stimulation pulses in relation to spontaneous (natural or intrinsic) cardiac activity. The sensed signals are referred to as an atrial electrogram or ventricular electrogram respectively. Note that the term electrogram lead is used herein to refer to the lead that transmits the sensed electrogram signal from the heart, and the term pacing lead is used to refer to the lead that transmits the stimulation pulse to the heart. As mentioned above, however, these "leads" are generally combined, i.e., the sensed electrogram signal is transmitted from the heart by the same lead that transmits the stimulation pulse to the heart. The separate terms "electrogram lead" and "pacing lead" are used herein merely to indicate that the electrogram signal and the stimulation pulse could be transmitted using separate leads. Every modern-day implantable pacemaker includes a sensing circuit, whether the activity of one or both chambers of the heart are sensed. The pacemaker sensitivity level is a measure of the gain of a sense amplifier. A cardiac event is sensed when an amplified electrogram signal exceeds a threshold value. If the sensitivity level is too low, i.e., the gain is too low, then some cardiac events will not be sensed because even peak signals may not exceed the threshold level. If the sensitivity level is too high, on the other hand, the high gain of the amplifier may cause noise signals to be sensed, giving rise to erroneous sensing of cardiac events. Pacemakers provided with communications telemetry (e.g.,noninvasive programming capabilities) advantageously allow the physician to set the sensitivity level. There are at least two disadvantages to having the physician set the sensitivity level. First, adjusting the sensitivity level is one more thing that the physician must remember to do, and it would be advantageous to relieve him or her of that task if it is possible to do so. Second, and more important, the physician generally sees the patient only occasionally, and weeks or months may go by without the sensitivity level being changed. Problematically, the sensitivity level that will accurately detect cardiac events at a given threshold level for a patient does not stay static. Changes in the sensitivity level are needed to accommodate for physical and mental stress. In addition, the sensitivity level needs to change as myocardial tissue (heart muscle tissue) undergoes scarring or other physical responses to the implanted electrogram lead(s). Other changes in the myocardium-electrogram lead interface, e.g., shifting of the position of the electrogram lead, may also cause changes in the proper sensitivity level. Unfortunately, these changes occur over a period of days and, in some cases, even hours or minutes. Because the physician generally sees the patient only every few weeks or months, the pacemaker sensing circuits can erroneously detect, or not detect, cardiac events over large periods of time. This erroneous detection/non-detection can cause under-pacing or over-pacing of the heart. Unfortunately for the patient, such changes may potentially leave him or her in a worse condition than he or she was in before the pacemaker was implanted. At best, the pacemaker is not able to operate efficiently --either by unnecessarily pacing and thereby draining the battery and risking pacemaker-induced tachycardias; or by not pacing as often as is needed by the patient. Thus, what is needed is a way to adjust the sensitivity level of a cardiac event detector in response to changing conditions in an electrogram signal over a short period of time. EP-A-0 469 817 relates to an arrhythmia control system employing arrhythmia recognition algorithm. One way of adjusting the sensitivity level of a cardiac event detector is discussed in U.S. Patent No. 4,708,144 issued to Hamilton et al. The Hamilton et al. patent shows the use of an attenuator to attenuate an amplified signal before such signal is digitized and rectified. After the signal is digitized and rectified, the signal is connected to a peak detector and to a digital comparator. The peak detector records the peak amplitude of the digitized and rectified signal and the digital comparator compares each digitized and rectified sample to a threshold value. If the digitized and rectified sample exceeds the threshold value, a cardiac event is detected. In response to the detected cardiac event, a pacemaker control circuit takes appropriate action. Each digitized and rectified sample is also presented to the peak detector which stores the maximum, or largest, digitized and rectified sample. The stored maximum sample is coupled to the pacemaker control circuit. After each cardiac event is sensed, the pacemaker control circuit averages the stored maximum with any of the previously occurring maximums yielding an average peak value. This average peak value is used to determine whether or not the attenuator should be adjusted to increase or decrease the attenuation provided by the attenuator. Specifically, if the average peak value increases, the attenuation is increased; and if the average peak value decreases, the attenuation is decreased, thereby adjusting the amplitude of the signal before it is digitized and rectified. Disadvantageously, even though the Hamilton et al. circuit provides one technique for dynamically adjusting the sensitivity level of a cardiac event detector, it lacks the ability to provide important cardiac monitoring functions beyond merely sensing a cardiac event. In addition to the detection of cardiac events, it is desirable, in the treatment of certain heart ailments, or for the detection of such ailments, to continuously monitor the patient over a certain period of time in order to determine the effectiveness of the treatment being administered by a pacemaker, under different conditions of stress or varying conditions of the heart. If the sensing circuit detects that the pacemaker is administering a less than ideal, or optimum, treatment, the treatment can be adjusted, e.g., by increasing or decreasing the rate at which the heart must beat before a pacing pulse is delivered, by decreasing the threshold level of the threshold detector, or by increasing the amplitude or duration of the pacing pulse. Unfortunately, events that would indicate that the pacemaker may be providing less than the optimum treatment may occur only infrequently. Thus, a physician may not detect such abnormal events during a weekly, biweekly or monthly examination, which may last only a few minutes and may not be able to adjust the pacemaker accordingly. In an effort to solve this problem, data acquisition systems have been developed that record electrogram signals over a predetermined period of time, e.g., on the order of days. The electrogram signals may then be analyzed by a physician or, in more advanced system, by a controller in the pacemaker in accordance with a control program that is designed to react to various conditions that are manifested by the electrogram signals. Such data acquisition systems advantageously allow detailed analysis of the electrogram signal over long periods of time thereby facilitating the detection and accommodation of infrequent heart abnormalities or the early detection of slowly developing heart ailments. Such long-term monitoring, particularly where implemented in advanced programmed systems, makes possible the purposeful and possibly automatic treatment of heart abnormalities long before the actual failure of a pacemaker to properly service the heart. In addition, with such automated systems; therapies, such as antitachycardia pacing and defibrillation, can be performed on the heart by pacing systems or dedicated defibrillators that would otherwise not be able to be performed as quickly or automatically. Unfortunately, implanted data acquisition systems have heretofore only been operable over a limited sample of the electrogram signal. This is because such systems store the electrogram signal in a memory. The memory is of a limited size, and when the memory is full, either part of the previously recorded electrogram signal must be discarded to make room for new electrogram signal to be recorded, or the data acquisition system must stop recording. In an effort to solve this problem, various high capacity means of storing electrogram signals have been developed such as magnetic tape recording systems. For example, U.S. Patent No. 4,250,888 issued to Grosskopf, suggests that when the memory is full, a warning message be given that alerts the patient to the need to contact the physician or to activate a tape recording system at home. Disadvantageously, the Grosskopf approach may require that the patient report to a potentially inconvenient location, i.e., the physician's office or the patient's home (where the tape recording system is located). Such inconvenience may encourage the patient to ignore the warning message. In addition, the warning message can be intrusive and embarrassing. Furthermore, such warning systems are not used with implantable pacemakers for at least two reasons. First, implantable pacemakers are implanted within the body and, as such, any warning means are neither visible nor readily heard. Second, implantable pacemakers must be compact and low power. Generally, the warning message is generated by a speaker or light source and thus draws a significant current. A problem with this prior art is therefore that the implantable event detection system is limited to operating on a small sample of the cardiac signal over a limited period of time, and that further it requires the use of inconvenient and impractical storage devices such as tape recording systems. Some systems have been developed that store only anomalous portions of the electrogram signal. See, e.g., Grosskopf . However, even these systems have a limited capacity and when a sufficient number of anomalous portions are stored, some data loss occurs. This data loss occurs when the memory is full and either the new signal must be discarded or the previously stored signal must be discarded. Problematically, the portion of the electrogram signal that is discarded may be the portion of signal that is needed for an accurate evaluation of the patient's heart condition. A problem with this prior art is that the cardiac event detection system is limited by the use of a finite capacity memory for storing the electrogram signal, and does not provide information sufficient for programmed evaluation in a controller and, if needed, automatic adjustment of a pacemaker or activation of a defibrillator in response to such evaluation. Another problem faced by the designers of automated cardiac pacing and/or defibrillation systems is the need for analysis of the electrogram signal. One approach to accurately analyzing the electrogram signal requires that the stored electrogram signal be subjected to complex digital filtering algorithms and statistical analysis. See e.g., U.S. Patent No. 4,422,459 issued to Simson. In order to generate the digitally filtered and statistically analyzed signals in Simson, a large computer system is employed. Such computer system is immobile and inconveniently located at, e.g., the physician's office, thus making implantation impossible. Disadvantageously, such algorithms and analysis require that many hundreds of mathematical operations be performed before an accurate conclusion as to whether the cardiac pacer and/or defibrillator are performing optimally can be obtained and, thus, before needed adjustment of the therapies provided by the cardiac pacer and/or defibrillator can be made. Problematically, this requires not only the use of a memory to store the incoming electrogram signal while the mathematical operations are being completed, the disadvantages of which are discussed above, but requires that many complicated computational steps be traversed by the controller. Such complicated computational steps are highly power-consuming --which would require more frequent replacement of the battery that powers the implantable cardiac pacer and/or defibrillator-- and thus, disadvantageous in implantable cardiac pacing applications. Another approach to accurately analyzing the electrogram signal has been to allow the physician to analyze the electrogram signal stored in a memory using conventional electrogram analysis techniques. Disadvantageously, in order to obtain the stored electrogram signal, the physician must download the stored electrogram signal via a telemetry circuit in the cardiac pacer system and/or defibrillator system. Hence, because a memory is used, the problems discussed above are also present in this approach. A further disadvantage of this approach is that no automated adjustment of the therapies provided by the cardiac pacer and/or defibrillator can be made because such adjustment must wait until the patient has travelled to the physician's office and until the physician has completed his or her analysis. Thus, what is needed is an implantable cardiac data acquisition and analysis system that does not require the use of complicated and highly power-consuming mathematical computations. The present invention advantageously addresses the above and other needs by providing an improved cardiac event detection system usable with implantable cardiac pacers (or pacemakers), defibrillators, or the like. According to the present invention there is provided an implantable cardiac data acquisition apparatus as claimed in the independent claim. Some optional features are set out in the claims depenent thereto. According to one embodiment, there is an implantable cardiac data acquisition system for detecting arrhythmic cardiac activity in a heart. The system is coupled to the heart via an electrogram lead as is known in the art of cardiac pacing. The system is also coupled to a therapy circuit that provides pacing and/or defibrillation therapies to the heart. An electrogram signal is sensed through the electrogram lead and is transmitted to a signal processor that is a part of the system. The signal processor preliminarily performs various signal processing functions on the electrogram signal that include filtering, amplifying and/or digitizing. The preliminarily processed and digitized electrogram signal is then coupled to an event detector and to a morphology detector. The event detector detects a cardiac event, e.g., an R-wave, within the electrogram signal whenever the electrogram signal exceeds a prescribed initial threshold value. Upon the detection of an event by the event detector, an event sensed signal is generated and coupled to a controller. The morphology detector detects various parameters associated with the morphology of the electrogram signal, and couples such parameters to the controller. (Note, as used herein the term "morphology" relates to the shape of the electrogram signal when viewed as a signal waveform as a function of time.) The controller generates and adjusts the threshold value used by the event detector, as well as control signals for controlling the operation of the signal processor as a function of the various morphology parameters sensed by the morphology detector. The controller further generates an output signal indicative of an arrhythmic cardiac condition in response to a prescribed change in said morphology parameters. Such output signal may then be used by the desired therapy circuit, e.g., a pacemaker or defibrillator, in order to deliver an appropriate therapy to the heart. The morphology detector in one embodiment includes one or more of the following: a minimum detector; a maximum detector; a peak detector; a baseline averager; a baseline sampler; an accumulator; and/or a counter; each described in more detail below. The minimum detector used in the morphology detector detects the magnitude of the most negative, i.e., below a baseline voltage, value in the electrogram signal and records the magnitude of such value as a minimum value. The maximum detector detects the magnitude of the most positive, i.e., above the baseline voltage, value in the electrogram signal and records the magnitude of such value as a maximum value. The minimum and maximum detectors thus generate a minimum signal and a maximum signal, respectively, that are indicative of the minimum and maximum values of the electrogram signal. The minimum and maximum signals are coupled to the controller. In another embodiment, the peak detector used in the morphology detector detects the largest value in the electrogram signal, regardless of whether the largest value is negative or positive. The peak detector records the magnitude of such largest value as a peak value and generates a peak signal that is indicative of the peak value. The peak detector is also coupled to the controller. In a further embodiment, the peak detector used in the morphology detector generates an average baseline value signal, indicative of the average magnitude of the electrogram signal over a first predetermined time period. Such average baseline value signal (hereafter "average baseline signal") is coupled to the controller. The first predetermined time period may be the time during which an average enable-signal is present, e.g., as determined by the counter or by the controller. As an alternative to the baseline averager, in accordance with one embodiment a baseline sampler may be used by the morphology detector. The baseline sampler generates an accumulated baseline signal, indicative of the accumulated magnitude (total baseline value) of the electrogram signal during the first predetermined period of time. The accumulated baseline signal is coupled to the controller. Typically, the baseline sampler operates digitally. That is, in one embodiment of the invention, the signal processor samples and digitizes the electrogram signal on a periodic basis so that the processed electrogram signal comprises a plurality of discrete (digital) sample values. The magnitude of each of the plurality of discrete sample values occurring during the first predetermined time period is added to the accumulated magnitudes of any discrete values that have occurred previously during the first predetermined time period. Thus, the total baseline value of the processed electrogram signal over the first predetermined time period is accumulated or totaled. In this way, the accumulated baseline signal is digitally generated. In accordance with another embodiment the signal processor does not digitize the electrogram signal, i.e. the processed electrogram signal is analog. The baseline sampler may rapidly sample the magnitudes of the processed electrogram signal over the first predetermined time period. The magnitude of each of the rapid samplings (discrete values) occurring during the first predetermined time period is added to the accumulated magnitude of any rapid samplings that have occurred previously during the first predetermined time period, thus, accumulating, or totaling, the accumulated magnitude of the rapid samples. In this way the accumulated baseline signal can be analogically generated. The accumulated baseline signal is coupled to the controller. Also during the first predetermined time period, in accordance with one embodiment the controller counts the number of values, i.e. the number of discrete values, in the processed electrogram signal that are accumulated by the baseline sampler. In this way a count value is generated. The total baseline value, from the baseline sampler, is then divided by such count value in order to generate the average baseline value. Alternatively, the average baseline value may be generated by generating a count signal indicative of the count value using the counter (as opposed to counting with the controller). Such count signal is then coupled to the controller, where the total baseline value from the baseline sampler is divided by such count value to generate the average baseline value. In a further embodiment, the accumulator, when used within the morphology detector, generates an accumulated cycle signal indicative of the accumulated magnitude of the electrogram signal during a cardiac cycle. The cardiac cycle begins, e.g., when the event is detected by the event detector as described above, and ends when a subsequent event is detected by the event detector in a similar manner. The accumulated cycle signal is then coupled to the controller. The accumulator may operate digitally or analogically in the same or similar manner as the baseline sampler described above. In addition to generating the average baseline value (which may be generated by the controller or the baseline averager as described above), the controller may generate one or more other output signals, such as signals indicating: a peak value; a fibrillation value; a morphology change value; and/or a gain change value as described below. A control program executed by the controller controls the generation of such output signals. The control program is executed, e.g., in response to the event sensed signal described above, and/or a time-out signal described below. The peak value output signal generated by the controller is the greater of the minimum value and the maximum value. (Note that the minimum and maximum values once determined, are both considered as positive magnitudes). Alternatively, the peak value may be generated by the peak detector as described above. The fibrillation value output signal generated by the controller is the peak value divided by the average baseline value. When the fibrillation value exceeds the value of a fibrillation threshold (stored, e.g., in the controller) the system sets a fibrillation flag. The fibrillation flag may then be used to engage a therapy circuit, e.g., a circuit that issues stimulation and/or defibrillation pulses in a prescribed pattern. The morphology change value output signal is generated by the controller in response to a change in the morphology of the electrogram signal as summarized below. In the discussion below, it is generally assumed that the maximum value is greater than the minimum value for the electrogram signal morphology being sensed; however, it is to be understood that in some instances the minimum value may be larger than the maximum value. Thus, while the following discussion is directed to setting the morphology change value based on changes in the maximum value, in the event that the minimum value is greater than the maximum value, the morphology change value would instead be set in response to changes in the minimum value. During normal sinus cardiac rhythm, the relationship between the maximum value and the minimum value does not typically change, i.e., the maximum value generally remains larger than the minimum value. In the event that the minimum value suddenly becomes larger than the maximum value, a radical change in the morphology of the electrogram signal would be indicated. The controller sets the morphology change value to "POLARITY" in response to such radical change. Furthermore, because it is likely that the signals thereafter generated by the morphology detector and/or the controller are no longer indicative of the morphology of the electrogram signal, the controller resets the morphology detector and/or the output signals generated by the controller in response to the morphology change value being set to "POLARITY." In addition to setting the morphology change value output signal to "POLARITY," the controller may set the morphology change value to "INCREASE," "DECREASE," and/or "NONE." For example, in the event that the present maximum value (or minimum value) is much greater than the average of the preceding maximum values (or minimum values), the controller sets the morphology change value to "INCREASE." When the morphology change value is set to "INCREASE," there is a high probability that a cardiac arrhythmia has begun. The therapy circuit may then respond to such detected arrhythmia as is known in the art of implantable pacemakers. In the event that the present maximum value (or minimum value) is much lower than the average of the preceding maximum values (or minimum values), but is still greater than the minimum value (or maximum value), the controller sets the morphology change value to "DECREASE." When the morphology change value is set to "DECREASE," there is a likewise high probability that a cardiac arrhythmia has begun, and the therapy circuit may thus respond accordingly. In the event that the present maximum value (or minimum value) is not much greater or much less than the average of the preceding maximum values (or minimum values), and the present maximum value (or minimum value) remains greater than the present minimum value (or maximum value), the morphology change value is set to "NONE." This indicates to the therapy circuit that the heart is experiencing normal sinus rhythm. In this way, changes in the morphology of the electrogram signal, particularly the onset of arrhythmias, can be detected without the need for complicated and power-consuming computation by the controller. Note that in the event the peak detector is used instead of the minimum and maximum detectors, the morphology change value is set in response to changes in the peak value in a manner similar to that described above. However, it is important to note that the morphology change value is generally not set to "POLARITY" when the peak detector is used instead of the minimum and maximum detectors. The gain change value output signal is generated by the controller in response to variations in an average of the previous maximum values (or minimum values). That is, in the event that the average of the previous maximum values (or minimum values) increases substantially, e.g., as such average is updated over a period of time, the controller sets the gain change value to "DECREASE" and decreases the gain of the sense amplifier (or other amplifier) used as a part of the signal processor. Similarly, if the average of the previous maximum values (or minimum values) decreases substantially, the controller sets the gain change value to "INCREASE" and increases the gain of the sense amplifier. If the average of the previous maximum values (or minimum values) does not change substantially, the gain change value is set to "NONE" and the gain of the amplifier is not changed. In this way, the gain of the sense amplifier, and thus the sensitivity of the cardiac event detection system, may be adjusted in response to varying amplitudes of the electrogram signal. In accordance with a further embodiment the controller sets the sensitivity of the event detector based, e.g., on the gain change value, or an average maximum or minimum value. In this way, the sensitivity of the event detector can be set to an optimum threshold value so that cardiac events are accurately sensed, while, e.g., electromagnetic noise and the like are not mistaken for cardiac events. The output signals of the controller representing the average baseline value, the peak value, the fibrillation value, the morphology change value, and the gain change value are used to generate the average baseline signal, a peak signal, a fibrillation signal, a morphology change signal, and a gain change signal, respectively. Such output signals are coupled to the appropriate therapy circuit, along with the fibrillation flag, and are used by the implanted therapy circuit to adjust the therapy services provided to the heart as needed. It is also thus apparent that such information is provided to the implanted therapy circuit without the need for complicated and highly power-consuming mathematical calculations in an implanted controller. In addition to the event detector and to the morphology detector, a time-out counter may be used with the controller. The time-out counter begins a time-out count in response to a time-out set signal. The time-out set signal is generated, e.g., by the controller in response to the event sensed signal. When the time-out count reaches a predetermined value, a time-out signal is generated and coupled to the controller. The time-out counter takes a predetermined period of time to reach the predetermined value after being set by the time-out set signal. If the time-out set signal is again received from the controller before the predetermined value is reached, no time-out signal is generated and the time-out detector begins the time-out count again. During normal sinus cardiac rhythm, the time-out detector receives the time-out set signal at a cardiac frequency with a period shorter than the predetermined period of time. Therefore, the time-out signal is not normally generated. However, if the cardiac rhythm deviates from normal such that the cardiac frequency has a period longer than the predetermined period of time, even for one period, the time-out signal is generated. The controller then executes the control program in response to the time-out signal. In this way the control program is executed when the cardiac frequency decreases below a predetermined frequency for even one period (in addition to being executed when the event sensed signal is generated as described above). Another aspect of the time-out detector of one embodiment is the ability to generate a cycle length signal indicative of the value of the time-out count held in the time-out counter when the next time-out set signal is received before the time-out count reaches the predetermined value. Such a cycle length signal is indicative of the period of the cardiac frequency whenever the period is shorter than the predetermined period of time. By dividing an accumulated cycle value (indicative of the accumulated cycle signal described above) by the cycle length signal, the controller generates a cycle average signal. The cycle average signal thus provides a measure of the average magnitude of the electrogram signal over the cardiac cycle. The method of acquiring information from an electrogram signal includes the steps of: (a) sensing an electrogram signal; (b) processing the electrogram signal to produce a plurality of discrete values, i.e., digitizing or sampling the electrogram signal; (c) recording at least one morphology value, e.g., the maximum value, the minimum value, the peak value, the accumulated cycle value, the average baseline value, the total baseline value, and/or the count value, associated with the plurality of discrete values, which morphology values are indicative of the morphology of the electrogram signal; and (d) generating at least one output signal indicative of any change in the morphology value; which at least one output signal provides information about the electrogram signal and may be used to control a therapy circuit. It is thus a feature to automatically adjust the sensitivity of a cardiac event detection system in response to varying electrogram signals. It is yet a further feature to provide such an implantable information acquiring system that also generates output signals that may be used by an implantable therapy circuit, such as a pacemaker, a defibrillator, or the like, and that facilitates automatic control and adjustment of the therapy provided by such therapy circuit. It is still an additional feature to provide such a system wherein the output signals used to control the therapy circuit may be generated without the need for performing complicated and highly power-consuming mathematical computations in an implanted controller. The above and other features will become more apparent from the Detailed Description of the Invention, presented in conjunction with the following drawings, wherein: FIG. 1 is a block diagram showing a cardiac event detector of the present invention in conjunction with an electronic pacing system; FIG. 2 is a basic block diagram of one embodiment of a cardiac event detector made in accordance with the present invention; FIG. 3 is a detailed block diagram of the cardiac event detector shown in FIG. 2; FIG. 4 is a graph showing an exemplary electrogram signal used as an input to the cardiac event detector of FIG. 2, and a digitized electrogram signal generated in response to the electrogram signal by a signal processor that is part of the cardiac event detector of FIG. 2; and FIGS. 5-1, 5-2, 5-3, 5-4 and 5-5 are a flowchart of an exemplary control program that may be used by a controller that is part of the cardiac event detector of FIG. 2. Like reference numerals are used to represent like elements in the various figures and the accompanying description that follows. The following description is of the best mode presently contemplated for carrying out the invention. This description is not to be taken in a limiting sense, but is made merely for the purpose of describing the general principles of the invention. The scope of the invention should be determined with reference to the claims. Referring to FIG. 1, a block diagram is shown of a cardiac detection system 21 used with an electronic pacing system 23. The cardiac detection system 21 and the pacing system 23 comprise, in combination, an implantable device 25 that is implanted into a patient and attached to a heart 10 of the patient via an electrogram lead 12. The electrogram lead provides electrical communication between the implantable device 25 and the heart 10. One signal that is communicated through the electrogram lead 12 from the heart 10 to the implantable device 25 is an electrogram signal. Such electrogram signals (sometimes referred to as ECG signals) are known in the art. See, e.g., U.S. Patent Nos. 4,596,255; 4,712,555; 4,817,605; and 4,940,052. The electrogram signal is coupled, within the implantable device 25, to the cardiac detection system 21. The implantable device 25 is housed in an implantable, hermetically sealed housing as is known in the art of implantable electronic pacemakers. The cardiac detection system 21 processes the electrogram signal and generates at least one output signal. The output signal is coupled to the pacing system 23, which pacing system may be considered as a therapy circuit. The therapy circuit 23 controls the therapy delivered to the heart 10 (typically stimulation pulses) via the electrogram lead 12. Alternatively, another lead, e.g., a pacing lead, may be used in lieu of or in addition to the lead 12 in order to deliver the desired therapy (e.g., stimulation pulses) to the patient's heart. By way of example, in FIG. 1, the therapy circuit 23 comprises a cardiac pacer. The cardiac pacer 23 includes a control circuit 20, a pulse generation circuit 22, a memory circuit 24, and a telemetry circuit 26. In practice, the control circuit 20 (which may hereafter be referred to simply as a "controller") and the memory circuit 24 (hereafter referred to as a "memory") may advantageously be utilized by both the cardiac detection system 21 and the cardiac pacer 23, thereby assuming both therapy-providing and event-detecting functions. However, to more clearly illustrate the present invention, the control circuit 20 and the memory 24 are shown as separate elements of the cardiac pacer 23, even though shared by the cardiac detection system 21. In some embodiments, the cardiac detection system 21 may utilize a separate controller from the control circuit 20 and memory 24 of the cardiac pacer 23. The control circuit 20 receives the output signal from the cardiac detection system 21 and, in response thereto, evaluates whether or not optimum therapy is being delivered to the heart 10. If the therapy being delivered is not optimum, the control circuit 20 makes adjustments, as required. In response to the adjustment of the therapy, the pulse generator 22 begins delivering a more optimum therapy to the heart 10. Several adjustments of the pulse generation circuit (or pulse generator) 22 may be needed before an optimum therapy is delivered. Under some circumstances, less than optimum therapy may be the best therapy that can be delivered. This is because the output signal may rapidly change in response to changing conditions of the heart 10 before needed adjustments can be made in response to previous output signals. However, the therapy delivered to the heart 10 will, in general, be repeatedly adjusted until optimum or near optimum therapy is delivered to the heart 10. In determining what is optimum therapy, based on the output signal generated by the detection system 21, the control circuit 20 may also use a memory 24. Various control parameters are stored in the memory 24 by a physician using a telemetry circuit 26. In order to store such parameters, the physician utilizes an external (non-implanted) programmer 28 that is coupled to the memory 24 and/or the control circuit 20 via the telemetry circuit 26 and a suitable communication link 27. Telemetry circuits used for this purpose are known in the art. Alternatively, various other therapy circuits 23 may be utilized with the cardiac detection system 21 of the present invention, such as other types of cardiac pacers or stimulators, implantable electrical defibrillators, implantable monitoring devices, and the like. Referring next to FIG. 2, a basic block diagram of the cardiac detection system 21 is shown. As seen in FIG. 2, the heart 10 and electrogram lead 12 are coupled to a signal processor 34. The signal processor 34, in turn, is coupled to a morphology detector 36 and an event detector 38. The two detectors 36 and 38 are coupled to a microcontroller 30. The microcontroller 30 is further coupled to the signal processor 34 and to the therapy circuit 23. As the electrogram signal is transmitted to the signal processor 34 via the electrogram lead 12, it is processed by the signal processor 34 into a plurality of discrete values. The discrete values are transmitted to the event detector 38 and to the morphology detector 36. The event detector 38 compares the amplitude of each of the discrete values with an event threshold value that is set by the microcontroller 30. In the event that the amplitude of one of the discrete values exceeds the event threshold value, an event sensed signal is generated by the event detector 38. The event sensed signal is then coupled to the microcontroller 30. Simultaneously, the morphology detector 36 evaluates each of the discrete values to determine the morphology of the electrogram signal. In response to this evaluation, the morphology detector 36 generates at least one morphology value that is coupled to the microcontroller 30. The morphology value is used by the microcontroller 30 to adjust the sensitivity of the signal processor 34, e.g., either by adjusting the gain of an amplifier or the attenuation provided by an attenuator or by adjusting the narrow band filter bandwidth. Additionally, the morphology value is used by the microcontroller 30 to generate an output signal. The output signal is coupled to the therapy circuit 23 over signal line 39. Such output signal adjusts the therapy circuit 23 to deliver a more optimum therapy to the heart 10 via the electrogram lead 12. Referring next to FIG. 3, a detailed block diagram of the cardiac detection system 21 is shown. The signal processor 34, event detector 38, morphology detector 36, microcontroller 30 and therapy circuit 23 are shown. In addition, a wide band filter 60 and an 8-bit analog-to-digital (A/D) converter 61 are shown coupled between the signal processor 34 and the microcontroller 30. Such wide band filter and A/D converter advantageously provide a signal path through which the microcontroller may directly monitor a digitized version of the electrogram signal in addition to the processing of the signal electrogram signal by the event detector 38 and the morphology detector 36. The signal processor 34 is comprised of a pre-amplifier 42 that is coupled to the electrogram lead 12, the wide band filter 60 and the microcontroller 30. A narrow band filter 44 is also part of the signal processor 34, and such filter is also coupled to the pre-amplifier 42. Finally, a 5-bit A/D converter 46 is a part of the signal processor 34. The 5-bit A/D 46 is coupled to the output of the narrow band filter 44. The output of the A/D 46 is directed to the morphology detector 36 and the event detector 38. The pre-amplifier 42 amplifies the electrogram signal from the electrogram lead 12 using a gain amplitude that is controlled by a 3-bit gain set signal from the microcontroller 30. The gain set signal is generated in response to a control program discussed more completely below in reference to FIGS 5-1 through 5-5. Included in the electrogram signal are various noise signals, e.g., undesirable components such as a T-wave, electromagnetic interference, myopotential voltage signals, and "baseline wander." (Baseline wander is a deviation, or drift, of the voltage of the electrogram signal during what should be the electrically quiet part of the electrogram signal.) Each of these noise components have been recognized in the art. After the electrogram signal is amplified, it is coupled to the narrow band filter 44. The narrow band filter 44 filters out most of the undesirable noise components in the amplified electrogram signal, and removes the need to compensate such signal for baseline wander in the control program. Sometimes it is desirable to widen or narrow the pass band of the narrow band filter 44, e.g., to adapt to different slew rates in the electrogram signal that occur at the onset of an arrhythmia, as well as to compensate for variations due to different electrogram lead types and positions. Advantageously, the pass bandwidth of the narrow band filter 44 can be adjusted in response to a high pass set signal obtained from the microcontroller 30. The high pass set signal is generated by the control program, or alternatively may be programmed to a desired value via the telemetry circuits and external programmer described above. It is known that arrhythmia signals have lower frequency content, so when the signal is not positively stable normal sinus rhythm, the bandwidth of the filter is widened. The output of the narrow band filter 44 is coupled to the 5-bit A/D converter 46. One type of 5-bit A/D converter 46 useable with the present invention transforms the amplified, filtered electrogram signal into one of thirty-two possible 5-bit digital codes. The conversion rate of the converter is about 1000 Hz, and the converter has two interleaved phases. During the first phase, a positive input produces a non-zero code (discrete value) indicative of the magnitude of the input, and a negative input produces a zero output. During the second phase, a positive input produces a zero output, and a negative input produces a non-zero output code (discrete value) indicative of the magnitude of the input. The 5-bit A/D converter 46 traverses one of these two phases during each conversion, and alternates between the first phase and the second phase during succeeding conversions. The 5-bit converter 46 also provides a polarity signal indicating the polarity of the last discrete value output therefrom, and an end-of-conversion signal indicating that an output code has been latched onto a 5-bit output bus of the converter 46. An exemplary 5-bit converter 46 that may be used with the present invention is available from Siemens Pacesetter, Inc. of Sylmar, California, which converter advantageously provides a low power consumption characteristic. The 5-bit output bus, polarity signal and the end-of-conversion signal are coupled to the event detector 38 and the morphology detector 36 over signal bus 37. As seen in FIG. 3, the event detector 38 comprises a threshold crossing detector circuit that generates an event sensed signal when an output from the 5-bit A/D converter 46 is larger than a threshold value. Note that this is an unsigned comparison: a discrete value of either polarity (negative or positive) that exceeds the threshold will cause the event sensed signal to be generated by the event detector 38. The microcontroller 30 sets a starting threshold value after the detection of a cardiac event, e.g., an R-wave. This starting threshold value is a function of the previous detected peak and average values of the R-wave as described more fully below in conjunction with FIGS. 5-2 and 5-3. Additionally, the microcontroller 30 prevents discrete values from being received by the threshold crossing detector (event detector 38) until a programmable, e.g., 0 to 50 ms, refractory time period has elapsed after the deliverance of the therapy to the heart 10 by the therapy circuit 23. Such action prevents spurious re-detections of, e.g., an R-wave, after the therapy has been delivered to the heart 10. One of the problems frequently encountered with some patients is a large T-wave, representing the repolarization of the cardiac tissue. The T-wave follows the depolarization of cardiac tissue, as manifested by the QRS complex, or R-wave. It is important in a cardiac event detection system that the T-wave not be confused with the R-wave. However, in patients exhibiting a large T-wave, the T-wave may be of the same order of magnitude as the R-wave. Hence, the detection circuitry, if detection is based solely on the electrogram signal exceeding a prescribed threshold value, has no way to distinguish T-waves from R-waves. The present invention advantageously addresses this problem by providing a threshold value to the event detector 38 that automatically adjusts itself to a low value (increased sensitivity) in order to best detect the R-wave, and (after the R-wave has been detected) to a high value (decreased sensitivity) in order to best avoid detection of the T-wave. After a sufficient time period at the decreased sensitivity, selected to avoid detection of the T-wave, the threshold value is then gradually decreased to its prior low value (i.e., the sensitivity is gradually increased) in anticipation of detecting the next R-wave. The manner of automatically adjusting the threshold value of the event detector 38 in accordance with the present invention is graphically depicted in FIG. 6. Represented in FIG. 6 are both an electrogram signal 600, containing an R-wave 604 and a T-wave 606, and a dynamic threshold value 602. Also included in FIG. 6 are "sample times," represented as tick marks along the horizontal axis. As described previously, the electrogram signal 600 is sampled each sample time by the A/D converter 46 (FIG. 3), with the resulting sampled value being examined by the threshold crossing detector 38. If at any given sample time the value of the electrogram signal is less than the threshold value, then no cardiac event detection occurs. If, however, at a given sample time the value of the electrogram signal is greater than the threshold value, then a cardiac event detection occurs. Thus, referring to FIG. 6, at sample time 614, before the occurrence of the R-wave 604, the threshold value is set to a low value (indicated as being 25% THRESHOLD) in anticipation of detecting an R-wave. At sample time 614, the electrogram signal is less than the threshold value so no R-wave detection has yet occurred. Between sample times 616 and 618, the R-wave 604 begins, causing the electrogram signal to cross over the threshold value, thereby signaling the occurrence of the R-wave. As soon as the R-wave is detected in this manner, i.e., at the next sample time 618 following the detection of the R-wave, the threshold value increases to its predetermined value (indicated as 100% THRESHOLD). Such predetermined threshold value is determined as a function of the detected peak and average values of prior cardiac events, as explained more fully below in conjunction with FIGS. 5-2 and 5-3. Alternatively, in some embodiments of the invention, the predetermined 100% THRESHOLD value may be a programmed value, selected by the physician at the time of implant or thereafter. Once the threshold value has been increased to the predetermined 100% THRESHOLD value (which threshold value represents a decreased sensitivity of the event detector 38), the threshold value remains fixed at such value for a predetermined period of time T F . Typically, T F is defined to be a programmable number of sample times of the A/D converter 46, e.g., 10 to 15 sample times (which corresponds to 10 to 15 ms, assuming a sample rate of 1000 Hz). The time T F during which the threshold value remains fixed at the 100% THRESHOLD value is selected to be sufficiently long to pass over a typical T-wave for the particular patient whose electrogram is being monitored. At the conclusion of the time T F , the threshold ramps down, in a controlled manner, to the 25% THRESHOLD value. In a preferred embodiment, such ramping down occurs by defining the threshold value to be a number of between 0 and 31, with 0 indicating the lowest (most sensitive) threshold value, and with 31 indicating the highest (least sensitive) threshold value. Thus, for example, if the 100% THRESHOLD value is determined to be 20, the 25% THRESHOLD value would be 5. As the threshold value ramps down from its 100% value to its 25% THRESHOLD value after the time period T F , it does so by decrementing the threshold value by one for a prescribed number of sample periods. For example, as shown by the ramp down curve 608 in FIG. 6, the threshold value is decremented by one for each two sample times, until the threshold value reaches the 25% THRESHOLD value. It some instances, it is desirable to adjust the ramp down rate of the threshold value. For example, while the ramp down curve 608 in FIG. 6 may be adequate for a given patient while at rest (having, e.g, a heart rate of 60 beats per minute, or one beat per second), it may be totally inadequate if the patient's heart rate suddenly increases because it may not be fully ramped down before the next R-wave occurs. Thus, the present invention includes a feature for automatically adjusting the ramp down rate as a function of the sensed heart rate. If the sensed heart rate is slow, then a relatively slow ramp down rate can be used, such as is represented by the ramp down curve 610 in FIG. 6 (decrementing one unit for each four sample times). If the sensed heart rate is moderate, then a moderate ramp down rate can be used, such as the ramp down curve 608 in FIG. 6 (decrementing one unit for each two sample times). If the sensed heart rate is fast, then a fast ramp down rate can be used, such as the ramp down curve 612 in FIG. 6 (decrementing one unit for each sample time). Advantageously, the adjustments of the threshold value, including the ramp down rate, may occur under control of the microcontroller 30 (FIG. 3) by simply providing the threshold crossing detector (event detector) 38 with a threshold set signal. The threshold set signal includes, in a preferred embodiment, at least three separate elements: a threshold set value; a threshold decrement interval value; and a detect refractory set value. The threshold set value may comprise, e.g., 5 bits (corresponding to 32 values). The threshold decrement interval value may comprise, e.g, 2 bits (corresponding to four values). Such decrement interval value effectively sets the ramp down rate, as described above, thereby providing four different ramp down rates. The detect refractory set value may comprise, e.g., 4 bits (corresponding to 16 values). The detect refractory set value effectively sets the duration of the time T F during which the threshold value remains fixed at the 100% THRESHOLD value. It is to be emphasized, of course, that the above values are only exemplary, and that any number of bits may be used to define the indicated variables in order to suit the needs of a particular application of the invention. As evident to those of skill in the art, the actual implementation of the threshold crossing detector 38 may occur in software, i.e., within the microcontroller 30 and/or in hardware, e.g., using appropriate registers and logic gates. Such registers and logic gates may be configured in conventional manner to digitally compare two digital numbers, one of which is the sample value and the other of which is the threshold value, and to decrement the threshold value as defined by the detect refractory set value (fixed threshold time, T F ) and the threshold decrement interval value (ramp down rate). Returning again to FIG. 3, it is seen that the morphology detector 36 is comprised of a plurality of detectors: a maximum detector 48, a minimum detector 50, an accumulator (or summer) 52, a baseline sampler 54, and a counter (or baseline timer) 56. Each of these detectors is coupled to the output bus 37 connected to the A/D converter 46. The maximum detector 48 latches the value on the output bus 37 if: (1) its magnitude exceeds the value currently stored by the maximum detector 48 (initially zero), and (2) the polarity output is positive, thereby always trapping the most positive discrete value. The minimum detector 50 latches the value on the output bus 37 if: (1) its magnitude exceeds the value currently stored by the minimum detector 50 (initially zero), and (2) the polarity output is negative, thereby always trapping the most negative discrete value. Thus, together the maximum and minimum detectors 48 and 50 completely bracket the excursion of the digitized electrogram signal. The maximum and minimum detectors 48 and 50 reset themselves, i.e., reset the value currently stored in the detectors 48 and 50 to zero, each time the microcontroller 30 reads them. This is normally at the end of a cardiac cycle. The accumulator 52 (labeled in FIG. 3 as an "unsigned summer") sums, or totals, all of the discrete values produced by the A/D converter 46 during the current cardiac cycle, i.e., since the previous event sensed signal was generated by the threshold crossing detector 38 without regard to polarity. Each time the A/D converter 46 issues an end-of-conversion signal from the end-of-conversion output, the conversion (discrete value) is latched and added to a previous accumulated cycle value in the accumulator 52, thereby forming a new accumulated cycle value. Thus, the accumulated cycle value is representative of the total magnitude of all of the discrete values issued by the A/D 46 during the current cardiac cycle. The baseline sampler 54 (labeled in FIG. 3 as an "unsigned baseline sampler") sums, or totals, all discrete values produced by the A/D converter 46 during a predetermined time period. The purpose of the baseline sampler 54 is to sample what should be the electrically quiet (baseline) part of the electrogram signal that lies between intracardial waveform complexes. During normal sinus cardiac rhythm, the sum collected by the baseline sampler 54 will be small and a sudden increase in the sum indicates the probability that a cardiac arrhythmia is in progress. The baseline sampler 54 is enabled by the baseline timer 56. The baseline timer 56 provides an enable signal that enables the baseline sampler during the predetermined period of time. The predetermined period of time is defined by two set time signals generated by the microcontroller 30: a set sample time signal and a set delay time signal. The set delay time signal determines a delay period after the detection of an event by the event detector 38 during which the enable signal is not present and, hence, during which the baseline sampler circuit is not enabled. The set sample time signal defines a sample period or sample window after the delay period during which the enable signal is present and hence during which the baseline sampler circuit 54 is monitoring the electrogram signal. After the sample period or window, the enable signal is not generated. By selecting appropriate values for the delay period, typically 100 to 150 ms, and for the sample period or window, typically 20 to 50 ms, the baseline sampler 54 is enabled during the quiet (baseline) part of the electrogram signal. While the baseline sampler 54 is enabled, it sums, or totals, all discrete values produced by the A/D converter 46. Each time the 5-bit A/D converter 46 issues an end-of-conversion signal from the end-of-conversion output, the conversion is latched and added to the previous total baseline value in the baseline sampler 54, thereby forming a new total baseline value. A time-out detector (cycle length counter) 58 is also shown in FIG. 3, and is coupled to the microcontroller 30. The cycle length counter 58 serves two functions. First, the cycle length counter 58 counts the number of end-of-conversion signals that are produced by the A/D converter 46 during a cardiac cycle, and thereby generates a cycle length value. (Note, each end-of-conversion signal occurs at a known rate, e.g., 1000 Hz, and thus has a known time period, e.g., 1 ms, associated therewith. Thus, the occurrence of 772 end-of-conversion signals, for example, indicates a time period of 772 ms. Each time the microcontroller 30 reads the cycle length counter 58, the cycle length counter 58 is cleared, i.e., the timer begins to count from zero. Second, the cycle length timer 58 generates a cycle time-out signal if the counted number of end-of-conversion signals exceeds a preset number. The preset number (which corresponds to a preset time period) is set by the microcontroller 30 in accordance with the control program or in accordance with a value stored in the memory 24 via the telemetry circuit 26 as described above (FIG. 1). The occurrence of the cycle time-out signal forces the microcontroller 30 to read the morphology detector 36 in the event no cardiac event is sensed during the preset time period (i.e., preset number of end-of-conversion signals). Referring next to FIG. 4, an exemplary electrogram signal 450 is shown to illustrate a typical input signal to the cardiac detection system 21 and to further graphically represent a digitized electrogram signal 451 generated in response to the electrogram signal 450 by the signal processor 34. Referring first to the electrogram signal 450, it can be seen that the electrogram signal is made up of five basic wave structures. First in the electrogram signal 450 is the Q-wave, immediately followed by the R-wave, and then the S-wave. These three waves are commonly referred to as the QRS complex. The QRS complex is followed by the T-wave, and then the barely discernable P-wave. After passing through the narrow band filter 44 (FIG. 3), the T-wave and the P-wave are usually not discernable. After the P-wave, the cycle repeats beginning with the QRS complex. The period T 3 of the electrogram signal, from QRS complex, to QRS complex, is referred to as the cardiac cycle. The period T 3 for the waveform 450 shown in FIG. 4 is approximately 800 to 300 ms (corresponding to a rate of approximately 70 to 200 ppm). The digitized electrogram signal 451, also shown in FIG. 4, reveals that the 5-bit A/D converter produces a non-zero output code 414 during a first sample period, approx. 1 ms, because the input electrogram signal 450 is positive when the electrogram signal is digitized and the A/D converter 46 is in the first phase, as mentioned above. During a second sample period, an output code 420 is produced that is zero because the input electrogram signal is still positive and the A/D converter 46 is in the second phase, as mentioned above. This process continues with every other sample period producing a non-zero output code (indicative of a positive input electrogram signal) until the occurrence of the Q-wave. With the occurrence of the Q-wave, a negative excursion of the electrogram signal first occurs, as indicated at 404. Immediately following the negative excursion 404, a positive excursion occurs at 401, representative of the positive R-wave being detected. The amplitude of the digitized electrogram signal at 401 exceeds the threshold value represented by line E-E. Exceeding the threshold E-E causes the event detector 38 to generate the event sensed signal described above. When the R-wave is detected, a delay period T 1 begins during which the baseline sampler 54 is disabled. Following the delay period T 1 , a sample period T 2 begins during which the baseline sampler 54 is enabled. As explained above, the baseline sampler is disabled after the sample period T 2 until the next sample period begins. The next delay period begins after the next event sensed signal is generated at 402. This will occur when the digitized electrogram signal again exceeds the threshold value designated by line E-E. It should be noted that the sample periods shown in FIG. 4 are, for reasons of clarity, not drawn to scale relative to the horizontal (time) axis. A typical cardiac cycle, for example, may be 1000 ms (corresponding to 60 heartbeats per minute). At a sample rate of 1000 Hz, one sample would be made every 1 ms, or 1000 samples would be made during a cardiac cycle. For purposes of clarity, however, only about 48 to 50 samples are shown in FIG. 4 as being taken during the cardiac cycle T 3 . Turning next to FIGS. 5-1, 5-2, 5-3, 5-4 and 5-5, a flowchart of a control program used by the microcontroller 30 is shown. Each main step of the flowchart is shown as a "box" or "block," with each box or block having a reference numeral associated therewith. The control program is called by the microcontroller whenever the microcontroller receives the event sensed signal from the event detector 38 or receives the cycle time-out signal from the time-out detector 58 (Block 1001). Under the direction of the control program, the microcontroller 30 reads the minimum value, the maximum value, the total baseline value and the cycle length value (Block 1003) of the electrogram signal during a given cardiac cycle. Note that the maximum and minimum detectors, the baseline sampler, and the time-out detector are reset (to zero) after they are read by the microcontroller. The average baseline value is then calculated by dividing the total baseline value by the number of sample delay values (Block 1005). Next the average baseline value is subtracted from the maximum value (Block 1007) and from the minimum value (Block 1009), thereby removing any normally distributed noise level. A variable called current polarity is set equal to the polarity value of the discrete value that caused the event sensed signal to be asserted. The following discussion assumes that the current polarity variable is positive, however, symmetrical operation would occur if the current polarity variable were negative. The current polarity is tested (Block 1011) and, if the polarity is positive, execution proceeds at Block 1013 (FIG. 5-2). The magnitude of the maximum value is compared with the minimum value (Block 1015). During normal sinus rhythm, the maximum value will be larger than the minimum value whenever the current polarity variable is positive. If the maximum value is smaller than the minimum value, a radical change in the morphology of the electrogram signal is indicated. In response to this radical change, the MORPH_CHANGE variable is set to "POLARITY" (Block 1029), the polarity variable is set to negative (Block 1031), the variable MIN_AVERAGE is set equal to the current minimum value (Block 1033) and the variable MAX_AVERAGE is set equal to the current maximum value (Block 1035). It is important that the MIN_AVERAGE and the MAX_AVERAGE variables be set to the current minimum and maximum values, respectively, because there is no reason to assume that these average variables are meaningful after such a radical change in morphology. The generation of MIN_AVERAGE and MAX_AVERAGE is discussed below. Execution of the program continues in Block 1037 (FIG. 5-3). If the maximum value is larger than the minimum value (as determined at Block 1015 of FIG. 5-2), the maximum value is compared to the MAX_AVERAGE variable (Blocks 1017 and 1019). If the maximum value is smaller than one-half times the MAX_AVERAGE variable, the variable MORPH_CHANGE is set to "DECREASE" (Block 1025); and if the maximum value is greater than one and one-half times the MAX_AVERAGE variable, the variable MORPH_CHANGE is set to "INCREASE" (Block 1027). If the maximum value is between one-half times the MAX_AVERAGE, and one and one-half times the MAX_AVERAGE, the variable MORPH_CHANGE is set to "NONE" (Block 1021). Next, the variable MAX_AVERAGE is updated using the following relationship: MAX_AVERAGE = ¾ (MAX_AVERAGE) x ¼ (maximum value) (Block 1023) and the threshold value is generated according to the following relationship: THRESHOLD VALUE = ½ (MAX_AVERAGE) + (average baseline value) (Block 1095). Note that the variable MIN_AVERAGE is similarly generated and used to generate the threshold value if the current polarity variable is negative (FIG. 5-3). Execution continues at Block 1097 (FIG. 5-4). The set THRESHOLD is then used by the event detector 38 as described previously. That is, the THRESHOLD value is written to the event detector as a threshold set signal (Block 1063, FIG. 5-1), and execution of the control program terminates (Block 1065) until another event sensed signal is coupled to the microcontroller or until a cycle time-out signal is received. Prior to writing the THRESHOLD value to the event detector, a fibrillation value is generated by dividing the maximum value by the baseline average value (Block 1101) (FIG. 5-4). The fibrillation value may be used, e.g., by a defibrillation program to set a fibrillation flag, or otherwise determine whether or not defibrillation therapy should be delivered to the heart. Next, the variable MAX_AVERAGE is tested. If MAX_AVERAGE is greater than twenty-four (Block 1103) then the gain of the amplifier is decreased by one-half (Block 1109) and the variable GAIN_CHANGE is set to "DECREASE" (Block 1115). If MAX_AVERAGE is less than eight (Block 1105) then the gain of the amplifier is doubled (Block 1111) and the variable GAIN_CHANGE is set to "INCREASE" (Block 1113). If the MAX_AVERAGE is between twenty-four and eight, the variable GAIN_CHANGE is set to "NONE" (Block 1051). Execution of the control program then continues at Block 1061 (FIG. 5-1). The GAIN_CHANGE variable is used by the pre-amplifier 42 to set its gain as described previously. As mentioned above, similar operation of the control program is exhibited in the event that the current polarity is negative. The primary difference in the execution of the control program when the current polarity is negative is that the minimum value is used to generate the variable MORPH_CHANGE and the threshold value (Blocks 1069 through 1091 of FIG. 5-3); and to generate the variable GAIN_CHANGE, and set the gain of the amplifier (Blocks 1045 through 1061 of FIG. 5-5). As described above, it is thus seen that the present invention provides a way of automatically adjusting the sensitivity (gain and/or threshold) of a cardiac event detection system in response to varying electrogram signals. It is further seen that the invention provides an implanted system that advantageously acquires information from an electrogram signal without the need far large memories to store all or portions of the electrogram signal. Additionally, it is see that the invention provides an implantable information acquiring system that generates output signals useable by an implantable therapy circuit, such as a pacemaker, a defibrillator, or the like, and that facilitates automatic control and adjustment of the therapy provided by such therapy circuit. Advantageously, it is seen that the output signals used to control the therapy circuit are readily generated without the need for performing complicated and highly power-consuming mathematical computations in an implanted controller. The method of detecting cardiac events occurring within a heart comprises the steps of: (a) sensing and transmitting an intrinsic electrogram signal of the heart to an implanted cardiac event detection apparatus; (b) processing the electrogram signal in the implanted cardiac event detection apparatus so as to convert it into a plurality of diserete values, each discrete value representing a sample of the electrogram signal at a specific sample time; and (c) examining the plurality of discrete values over time within said implanted cardiac event detection apparatus to determine at least one morphology value indicative of variations in the morphology of the electrogram signal. Preferably, step (b) includes (b1) filtering said electrogram signal using a band-pass filter; (b2) amplifying said filtered electrogram signal; and (b3) converting said amplified filtered electrogram signal from an analog signal to a digital signal, said digital signal comprising said plurality of discrete values, each of said plurality of discrete values representing the magnitude of said electrogram signal at a specific sample time. Preferably, step (c) (c1) recording a minimum value of said plurality of discrete values; and (c2) recording a maximum value of said plurality of discrete values; said minimum and maximum values providing an indication of one type of variation in the morphology of said electrogram signal. Step (c) may also include (c3) accumulating a total baseline of a first portion of said plurality of discrete values; said total baseline value providing an indication of another type of variation in the morphology of said electrogram signal. Alternatively, step (c) may include: (c1) examining said plurality of discrete values to determine a peak value; and (c2) accumulating a total baseline value of a first portion of said plurality of discrete values; said peak value and total baseline value providing an indication of variations in the morphology of said electrogram signal, which variations in turn provide an indication of the occurrence of a cardiac event.

Claims: An implantable cardiac data acquisition apparatus (25) for detecting arrhythmic cardiac activity in the heart (10) comprising: an electrogram lead (12) coupled to said heart that transmits an electrogram signal from said heart; a signal processor (34) coupled to the electrogram lead that receives the electrogram signal and digitises it into a plurality of discrete values, said signal processor including: a peak detector (36) for detecting at least one member in a peak group, said peak group comprising a minimum value in said plurality of discrete values, a maximum value in said plurality of discrete values, and a peak value in said plurality of discrete values; and an event detector (38) coupled to the signal processor that detects the occurrence of a cardiac event in the electrogram signal; characterised by a baseline sampler for accumulating a total baseline value of a first portion of said plurality of discrete values; controller means (20) for counting the number of values in the first portion of said plurality of discrete values, and means for dividing said total baseline value by the number of counted values, thereby calculating an average baseline value for the first portion of said plurality of discrete values; and said controller means (20) including means (22) for generating a fibrillation signal that comprises the quotient of said peak value divided by said average baseline value, a change of a prescribed amount in said fibrillation signal providing an indication of the presence of arrhythmic cardiac activity in the heart; wherein said event detector (38) detects the occurrence of a cardiac event in the electrogram signal whenever the magnitude of said electrogram signal exceeds a threshold value, said threshold value comprising one of the control signals generated by said controller means (20); and wherein said controller means (20) includes means for automatically changing the threshold value from an initial value to a temporary value in response to the detection of a cardiac event, and then gradually returning the threshold value from the temporary value back to the initial value within a prescribed time period, said prescribed time period being selected to return the threshold value to its initial value prior to the occurrence of the next cardiac event. The implantable apparatus, as set forth in claim 1, wherein said controller means (20) automatically sets said initial threshold value as a function of the previous peak values determined by said peak detector (36). The implantable apparatus, as set forth in claim 2, wherein said controller means (20) sets said temporary threshold value to be a prescribed percent of said initial threshold value. The implantable apparatus, as set forth in any of claims 1 to 3, wherein said automatic threshold changing means of said controller means (20) includes means for maintaining the threshold value at said temporary value for a programmed period of time, and then ramping the threshold value from its temporary value back to its initial value in accordance with a programmed slope. The implantable apparatus, as set forth in claims 1 to 4, wherein said controller means (20) further provides control signals that control the operation of said signal processor (34). An apparatus as claimed in any preceding claim comprising: a morphology detector (36) comprising said peak detector coupled to the signal processing means that detects changes in the morphology of the electrogram signal and generates at least one morphology value signal indicative of the morphology of the electrogram signal; said controller means (20) being coupled to the morphology (36) and event detectors (38) for executing a control program, said control program being initiated upon the occurrence of said cardiac event and using said at least one morphology value to generate at least one output signal comprising said fibrillation signal. The apparatus of claim 6 wherein said morphology detector (36) includes: a minimum detector that detects a minimum value in said plurality of discrete values; and a maximum detector that detects a maximum value in said plurality of discrete values; said minimum and maximum values providing an indication of one type of change in said electrogram signal. The apparatus of claim 7 wherein said total baseline value provides an indication of another type of change in said electrogram signal. The apparatus of claim 8, wherein said controller means (20) generates a peak value that is equal to the largest absolute magnitude of said minimum value detected by said minimum detector and said maximum value detected by said maximum detector, said peak value providing an indication of the change, if any, in the morphology of the electrogram signal from cardiac event to cardiac event, thereby providing an indication of whether arrhythmic cardiac activity is present. The apparatus of claim 9, wherein said fibrillation signal provides an indication that said heart may be experiencing fibrillation. The apparatus of claim 1, wherein said signal processor means (34) comprises: a filter stage coupled to said electrogram lead that filters said electrogram signal; an amplifier stage (42) coupled to said filter stage that amplifies the electrogram signal filtered by said filtered stage; and an analogue-to-digital (A/D) processor stage (46) coupled to said amplifier stage, said A/D processor stage being configured to sample the filtered electrogram signal provided by said amplifier stage at a prescribed sampling rate and convert each sample of the filtered electrogram signal to a digital signal, the digital signals thus generated comprising said plurality of discrete values. The apparatus of claim 11, wherein said amplifier comprises a linear amplifier that amplifies the electrogram signal by a set gain factor, and wherein said controller includes means for generating said set gain factor as a function of said at least one morphology value.

Inventors: BORNZIN GENE A, MCCLURE KELLY H, BORNZIN, GENE A., MCCLURE, KELLY H.
Applicants: PACESETTER INC, PACESETTER, INC.
Application year: 1993
Main classification: A61N1
IPC codes: A61B5, A61N1