Heart simulator

A development system for a cardiac pacemaker control system, the development system comprising: a cardiac pacemaker prototype including an adaptive control system; and a heart simulator in communication with the cardiac pacemaker prototype comprising: a means for receiving pacing stimulations output from the cardiac pacemaker prototype; and a hemodynamic sensor model responsive to the received pacing stimulations, the hemodynamic sensor model being operable to output a signal simulating a hemodynamic sensor in a plurality of heart states, the hemodynamic sensor model being further operative to output the signal simulating the hemodynamic sensor in the plurality of heart states in a plurality of heart conditions.

BACKGROUND OF THE INVENTION

The invention relates generally to the field of heart simulators, and in particular to a simulator providing temporally dependent waveforms simulating implanted hemodynamic sensor signals.

Cardiac resynchronization therapy (CRT) is an established therapy for patients with congestive systolic heart failure and intraventricular electrical or mechanical conduction delays. CRT is based on synchronized pacing of the two ventricles according to the sensed natural atrium signal that determines the heart rhythm. The resynchronization task demands exact timing of the heart chambers so that the overall stroke volume is maximized for any given heart rate (HR). Optimal timing of activation of the two ventricles is one of the key factors in determining cardiac output. The two major timing parameters which are programmable in a CRT device and determine the pacing intervals are the atrioventricular (AV) delay and interventricular (VV) interval. The AV delay is defined herein as the delay from an atrial event that triggers filling of the atrium to a ventricular event that triggers blood ejection from the ventricle. The VV interval is defined herein as the interval between the pacing signal for the left ventricle to the pacing signal for the right ventricle.

Rate responsive pacemakers were designed in the 1980's and 1990's that allow the atrial pacing rate to automatically change according to combined inputs from two implanted sensors typically, accelerometer and minute ventilation sensors. The rate responsive system give a more physiologic pacing to Bradycardia patients and allow higher cardiac output at exercise conditions.

Zachary I. Whinnett et al in “Haemodynamic effects of changes in AV and VV delay in cardiac Resynchronization Therapy show a consistent pattern: analysis of shape, magnitude and relative importance of AV and VV delay”, Heart published online, 18 May 2006, doi:10.1136/hrt.2005.080721, studied the importance of the AV delay and VV interval optimization in CHF patients. The authors conclude that changing the AV and VV delay result in a curvilinear and reproducible acute blood pressure response. This shape fits very closely to a parabola which may be helpful in designing a streamlined clinical protocol to select optimal AV and VV delay.

Various types of implantable electrodes are currently available which generate intra-cardiac electrograms (IEGMs). These electrodes are bi-directional and are capable of both outputting electrical signals reflecting heart electrical activity and receiving electrical impulses to affect heart activity. Various types of hemodynamic sensors are also available reflecting heart mechanical behavior, including without limitation, impedance sensors, pressure sensors and cardiac wall accelerometers, QT interval sensors and non-invasive hemodynamic sensors. However, there are currently no commercially available CRT devices that modify the AV delay and VV interval responsive to the implanted hemodynamic sensors in an adaptive, closed loop system. The ability to design and validate such system is a major technological hurdle for developing and bringing to market closed loop physiological pacemakers and defibrillators that will deliver optimal therapy to CHF and other heart disease patients. In particular the design and validation of the control system which processes the data and makes the appropriate adaptation and classification of patient condition on-line would be greatly enhanced by an appropriate simulator.

Cardiac Output (CO) is defined as the blood volume in milliliters pumped by the heart in one minute. HR and stroke volume, defined as milliliters of blood pumped/systole, determine CO. CO is influenced by the autonomic nervous system in that 1) sympathetic stimulation increases heart rate and stroke volume and 2) parasympathetic stimulation slows heart rate. CO is also influenced by increased venous return, as would occur during exercise. The increased venous return causes greater stretch of the cardiac fibers resulting in greater contraction (i.e., increased stroke volume). This increase in stroke volume due to increased venous return is called Frank-Starling law of the heart.

Contractility of the heart is a major determinant of its ability to pump blood around the body, which is required for the functions of all organs. Because the heart is a muscle, it can adapt to different conditions/requirements that directly affect its contractility and efficiency.

U.S. Pat. No. 3,833,865 to Palmer, issued Sep. 3, 1974, the entire contents of which is incorporated herein by reference, is addressed to a heart simulator which generates an electric waveform similar to that produced by the human heart. Unfortunately such a device is neither responsive to the output of an adaptive CRT device, nor is their any provision for hemodynamic sensor simulation.

U.S. Provisional patent application Ser. No. 60/656,392 filed Feb. 28, 2005 to Rom, entitled “Adaptive Cardiac Resynchronization Therapy and Vagal Stimulation System”, the entire contents of which is incorporated herein by reference, is addressed to a method an apparatus for optimizing cardiac resynchronization therapy devices and vagal stimulators.

U.S. Provisional patent application Ser. No. 60/685,464 filed May 27, 2005 to Rom, entitled “Ventricle Pacing During Atrial Fibrillation Episodes”, the entire contents of which is incorporated herein by reference, is addressed to a system that learns to associate ventricle-atrial intervals with temporal patterns of at least one hemodynamic sensor using neural network processing.

U.S. Provisional patent application Ser. No. 60/897,513 to Rom, entitled “Intelligent Control System for Adaptive Cardiac Resynchronization Therapy Devices” filed Jul. 17, 2006 to Rom, the entire contents of which is incorporated herein by reference, is addressed to an adaptive CRT control system that achieves optimal AV delay and VV pacing intervals with temporal patterns of stroke volume comprising a learning module and an algorithmic controller module supervising the learning module.

International patent application S/N PCT/IL2004/000659 published Jan. 25, 2005 as WO 2005/007075 to Rom, entitled “Adaptive Resynchronization Therapy System”, and filed as U.S. patent application Ser. No. 10/565,279 Jan. 20, 2006, the entire contents of each of which is incorporated herein by reference, is addressed to a system including a learning module and an algorithmic module for learning a physiological aspect of a patient body and regulating the delivery of a physiological agent to the body. Such an adaptive system is advantageous however its development and validation would be simplified by the existence of a heart simulator providing signals simulating hemodynamic sensors.

Thus there is a need for a heart simulator that generates simulated IEGMs and simulated hemodynamic sensor outputs. Preferably the simulated outputs would be responsive to an adaptive CRT device. Such a heart simulator would be particularly useful for the development of closed loop pacemaker and defibrillator controllers working according to implanted hemodynamic sensors thereby shortening the lead time before in-vivo pre-clinical and clinical experiments can begin.

SUMMARY OF THE INVENTION

Accordingly, it is a principal object of the present invention to overcome the disadvantages of prior art heart simulators. This is provided for in the present invention by a heart simulator that generates simulated implanted IEGM signals and hemodynamic sensor signals thereby simulating both heart electrical and mechanical activity, the simulation signals being responsive to the output of a CRT device. Preferably, the heart simulator further exhibits a programming input, allowing for simulation of various heart action irregularities.

The invention provides for a development system for a cardiac pacemaker control system, the development system comprising: a cardiac pacemaker prototype including an adaptive control system; and a heart simulator in communication with the cardiac pacemaker prototype comprising: a means for receiving pacing stimulations output from the cardiac pacemaker prototype; and a hemodynamic sensor model responsive to the received pacing stimulations, the hemodynamic sensor model being operable to output a signal simulating a hemodynamic sensor in a plurality of heart states, the hemodynamic sensor model being further operative to output the signal simulating the hemodynamic sensor in the plurality of heart states in a plurality of heart conditions.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

The present embodiments enable a heart simulator that generates simulated implanted IEGM signals and hemodynamic sensor signals thereby simulating both heart electrical and mechanical activity, the simulation signals being responsive to the output of a CRT device. Preferably, the heart simulator further exhibits a programming input, allowing for simulation of various heart action irregularities.

The aim of the heart simulator is to allow the design and testing of a prototype pacemaker and hence it is advantageous that the heart simulator be responsive to outputs of the prototype pacemaker and be versatile enough to simulate various physiological phenomenon that may occur in a real heart and be close enough to the appropriate physiological signal to aid in pacemaker development. The sensor model does not have to mimic exact patient behavior and exact specific values, particularly if used in the development of a pacemaker comprising a neural network which learns from the response of the attached heart.

FIG. 1illustrates a high level block diagram of a system10for design verification of a cardiac pacemaker comprising a heart simulator20, a cardiac pacemaker prototyping board30and a computing device40in accordance with a principle of the current invention. Heart simulator20outputs a simulated hemodynamic sensor signal50for input to cardiac pacemaker prototyping board30and further exhibits a bidirectional IEGM path60which carries both a simulated IEGM70from heart simulator20to cardiac pacemaker prototyping board30, preferably comprising an cardiac pacemaker in development having an adaptive control system, and a pacing stimulation80generated by cardiac pacemaker prototyping board30to heart simulator20. In an exemplary embodiment simulated IEGM70output by heart simulator20and simulated hemodynamic sensor signal50are responsive to pacing stimulation80received by heart simulator20from cardiac pacemaker prototyping board30. Computing device40is connected by a bidirectional data path90to heart simulator20and by a bidirectional data path95to cardiac pacemaker prototyping board30. Simulated IEGM70, pacing stimulation80and simulated hemodynamic sensor signal50may each represent a plurality of signals without exceeding the scope of the invention.

In operation, cardiac pacemaker prototyping board30receives simulated IEGM70and simulated hemodynamic sensor signal50from heart simulator20. In an exemplary embodiment in which cardiac pacemaker prototyping board30comprises a learning module as described in international patent application S/N PCT/IL2004/000659 published Jan. 25, 2005 as WO 2005/007075 to Rom, entitled “Adaptive Resynchronization Therapy System” referenced above, cardiac pacemaker prototyping board30is operative to learn the heart condition as indicated by the received simulated IEGM70and simulated hemodynamic sensor signal50and in response deliver appropriate pacing stimulation80to heart simulator20through bidirectional IEGM path60.

Heart simulator20is programmable to a predetermined condition by computing device40via bidirectional data path90, and in response outputs signals representative of the predetermined electrical heart condition via simulated IEGM70and the predetermined mechanical heart condition via simulated hemodynamic sensor signal50. Heart simulator20is further responsive to pacing stimulation output by cardiac pacemaker board30via pacing stimulation80to change the simulated heart condition as indicated by simulated IEGM70and simulated hemodynamic sensor signal50.

Computing device40enables the presentation in real time of data from both heart simulator20and cardiac pacemaker prototyping board30and can also run test codes on heart simulator20and test the design performance of cardiac pacemaker prototyping board30. In one embodiment, cardiac pacemaker prototyping board30is provided with a wireless communication device for wireless data transfer to computing device40, and thus bidirectional data path95represents a wireless path. Data accumulated in cardiac pacemaker prototyping board30is transferred to computing device40for performance monitoring via bidirectional data path95.

FIG. 2illustrates a high level block diagram of an embodiment of cardiac pacemaker prototyping board30ofFIG. 1comprising: an analog interface sensing and pulse generator110; a field programmable gate array (FPGA)120; an input/output device130; a bidirectional data connection95; a bidirectional IEGM path60carrying both a simulated IEGM70and a pacing stimulation80; and a simulated hemodynamic sensor signal50. FPGA120comprises the digital design of the pacemaker device control system and is connected to both input/output device130and analog interface sensing and pulse generator110. FPGA120can implement for example both an adaptive CRT device learning module, a spiking neural network, and a deterministic controller of the pacemaker device as described in international patent application S/N PCT/IL2004/000659 published Jan. 25, 2005 as WO 2005/007075 entitled “Adaptive Resynchronization Therapy System” to Rom referenced above. Analog interface sensing and pulse generator110comprises an analog to digital interface that senses the electrical signals representative of the simulated electrical heart condition received via simulated IEGM70and the simulated mechanical heart condition via simulated hemodynamic sensor signal50, amplifies and digitizes the incoming signals. Analog interface sensing and pulse generator110further comprises a pulse generator that stimulates heart simulator20through pacing stimulation80.

Bidirectional data connection95may comprise a cable connection that allows for programming FPGA120from an external device via input/output device130to which it is connected. Bidirectional data connection95may further comprise a wireless data connection for uploading data accumulated in cardiac pacemaker prototyping board30and to receive programming instructions.

FIG. 3illustrates a high level block diagram of a heart simulator20in accordance with a principle of the current invention comprising a hemodynamic sensor model200, an electrical heart model210, and input/output interface220, a simulated hemodynamic sensor signal50, a simulated IEGM70, a pacing stimulation80and a bidirectional data path90. Hemodynamic sensor model200comprises an AV timer230, a finite state machine (FSM)235, a QT timer240and an optimal timer245. Electrical heart model210is operational to generate simulated IEGM80representative of a particular heart condition, and in particular the electrical behavior thereof. Hemodynamic sensor model200is operational to generate simulated hemodynamic sensor signal50representative of a particular heart condition, and in particular the mechanical behavior thereof as sensed by a particular specific implanted sensor. Bidirectional data from bidirectional data path90is connected to input/output interface220and input/output interface220is connected to each of hemodynamic sensor model200and electrical heart model210. Electrical heart model210is in communication with hemodynamic sensor model200.

Advantageously, various types and designs of hemodynamic sensor are modeled by hemodynamic sensor model200allowing for testing of cardiac pacemaker prototyping board30ofFIG. 1in combination with different types of sensors and sensor features. In one embodiment hemodynamic sensor model200comprises a model of one or more impedance sensors, in another embodiment hemodynamic sensor model200comprises a model of one or more pressure sensors, in another embodiment hemodynamic sensor model200comprises a model of one or more cardiac wall accelerometers, in another embodiment hemodynamic sensor model200comprises a model of one or more QT interval sensors, and in another embodiment hemodynamic sensor model200comprises a model of one or more non-invasive hemodynamic sensors. In another embodiment hemodynamic sensor model comprises a plurality of sensor types selected from impedance sensors, pressure sensors and accelerometers. Thus, different types of sensors and sensor features, in combination with different cardiac pacemaker architectures, optimization parameters and learning methods can be tested and verified using heart simulator20, and only after satisfactory results are obtained, an application specific integrated circuit (ASIC) can be fabricated and tested in an implant device in animal and clinical trials which are both an orders of magnitude more time consuming and expensive.

There are several implanted hemodynamic sensors that are being developed by manufactures of pacemakers and others which are primarily targeted at closed loop implanted cardiac pacemaker devices. These hemodynamic sensors include, but are not limited to, ventricle impedance sensors, ventricle pressure sensors, cardiac wall acceleration sensors and QT interval sensors.

The present invention does not require modeling in detail the complete mechanical behavior of the heart in order to generate hemodynamic sensor model200. Thus, there is no requirement for an electromechanical heart model that models each heart chamber behavior, each heart valve and cell conductance. Instead, hemodynamic sensor model200of the subject invention focuses only on modeling the specific sensor signal temporal patterns as a function of the heart condition and pacing interval. It is this information which will be sensed by the implanted device through the specific implanted sensors, and cardiac pacemaker prototyping board30is responsive to this input data. In a preferred embodiment hemodynamic sensor model200comprises a plurality of models, including a model for each type of sensor that will be used by cardiac pacemaker prototyping board30.

The invention will hereinafter be described in detail in relation to an impedance sensor model, and in particular to a ventricle impedance sensor model, however this is by way of illustration only and is not meant to be limiting in any way. The invention is equally applicable to pressure sensors, accelerometers and interval sensory.

Hemodynamic sensor model200generates a temporal impedance signal. Stroke volume and pre-load volume measures can be extracted from the temporal impedance signal beat after beat. Hemodynamic sensor model200of the present invention preferably uses a state machine description of the sensor output with a pre-defined functional dependence of the main signal characteristics as will be explained below, however this is not meant to be limiting in any way. The use of fuzzy logic or parallel processing is specifically contemplated. AV timer230, QT timer240, FSM235and optimal timer245are each operational, as will be described further hereinto below, to generate sub-optimal states representative of sub-optimal pacing.

FIG. 4illustrates a high level state machine representation of hemodynamic sensor model200of heart simulator20in accordance with a principle of the current invention, and in an exemplary embodiment is implemented by FSM235ofFIG. 3. FSM235exhibits5states, state300which simulates passive filling, state310which simulates active filling, state320which simulates ejection, state330which simulates regurgitation and state340which simulates late passive filing. The output of hemodynamic sensor200ofFIG. 3is a function of: the state in accordance with pre-defined formulas that describe the dependence of the impedance signals on various pre-defined functions; user set parameters such as heart rate, ventricle volume pre-load, optimal AV delay and QT delay that depend on heart rate; and the timing information of electrical heart model210such as a sensed atrial event and a sensed or paced ventricle event. Advantageously, electrical heart model210and hemodynamic sensor model200are responsive to pacing stimulation80.

FSM235of hemodynamic sensor model200begins at passive filling state300and the impedance value output by hemodynamic sensor model200decreases with time in this state. It is to be understood that the impedance value decreases with time in passive filling state300since the ventricle is filling with blood and the heart impedance has an inverse relation to ventricle volume. When an atrial event350occurs, FSM235sets AV timer230ofFIG. 3, as illustrated at state360, with the optimal AV delay, as will be explained further hereinto below. In state365an optimal timer is loaded with the optimal AV delay plus the QT interval, as will be explained further hereinto below and makes a transition to active filling state310.

Active filling state310is characterized by a decreasing impedance at an increased slope due to the fact that blood is drawn into the ventricle of the heart at an increased rate due to contraction of the atrium.

Active filling state310continues until either a paced ventricular event occurs, denoted as event370, as determined by pacing stimulation80, or AV timer230set by event360expires as denoted by event380. In the event ventricular pacing event370occurs, FSM235sets QT timer240ofFIG. 3, as illustrated at state390, with a pre-determined QT interval, as will be explained further hereinto below, and transitions to ejection state320, as optimal behavior. In ejection state320the impedance value output by hemodynamic sensor model200begins to increase since in ejection state320the blood is ejected from the ventricle and the volume decreases resulting in increased impedance. In the event QT timer240ofFIG. 3, loaded in state390with the pre-determined QT interval, expires, FSM235of hemodynamic sensor model200resets the amplitude in accordance with Eq. 2 described below as denoted by event410and transitions to passive filling state300as described above, thereby restarting the cycle, as optimal behavior.

In the event that in active filling state310, AV timer230set in event360expires, i.e. the optimal AV delay has expired and a paced ventricular event has not occurred, FSM235of hemodynamic sensor model200transitions to regurgitation state330which is representative of backflow to the atrium as illustrated by event380, as suboptimal behavior. In regurgitation state330the impedance increases since blood is backflowing to the atria resulting in a lower ventricular volume. Regurgitation state330is maintained until a paced ventricular event occurs, as denoted by event420, and in response FSM235of hemodynamic sensor model200sets QT timer240ofFIG. 3, as illustrated at event430, with a pre-determined QT interval and transitions to ejection state320.

In the event that in ejection state320optimal timer245set in state365expires, denoted event450and characteristic of ejection state320being entered late due to having experienced regurgitation state330, FSM235of hemodynamic sensor model200transitions to late passive state340. In late passive state340the impedance value output by hemodynamic sensor model200increases over time representative of late passive ejection of blood, however due to the delayed paced event, i.e since the AV timer230expired as described above in relation to state380, ejection began too late and thus the impedance waveform will not reach the expected maximal amplitude in this cardiac cycle. This is representative of suboptimal ejection.

In the event that in late passive state340a right atrial event occurs, denoted event470, FSM235transitions to active filling state310and sets both AV timer230and optimal timer245as described in relation to states360,365. In the event that in late passive state340QT timer240set in state430expires, denoted event460, FSM235transitions to passive filling state300.

Both regurgitation state330and late passive state340exhibit a reduced stroke volume as compared with optimal behavior, which exhibits an optimal stroke volume. As indicated above, stroke volume can be extracted from the temporal impedance signal. The stroke volume is represented by the difference between the maximum and minimum impedance amplitude as will be described further hereinto below.

As described above in relation toFIGS. 3 and 4there are a few pre-defined characteristic functions and timers that determine the simulated temporal impedance signal waveform.

AV Timer

AV timer230ofFIG. 3is used to transition from active filling state310to regurgitation state330when the pre-defined optimal AV delay expired before a ventricular pacing event occurs. The pre-defined optimal AV delay depends on the cardiac cycle length set by the user, preferably utilizing computing device40, through a simplified formula:
AVopt=160*Cardiac Cycle/1000 (msec)  Eq. 1
with the 160 representing a typical value for a normal AV interval at rest. The optimal AV delay, AVopt, and any deviation from pacing in accordance with AVoptinfluences the simulated impedance waveform amplitude. The larger the deviation from AVoptthe larger the deviation from the optimal impedance values as will be described further hereinto below.
QT Timer

QT timer240is used to transition from ejection state320to passive filling state300when QT timer240expires, at the pre-defined QT interval as shown by event400, or to transition from late passive state340to passive filing state300, as shown by event460, if the pacing ventricular event420occurred late after the AV timer expired in regurgitation state330. The pre-defined QT interval setting for QT timer240is a function of the user set cardiac cycle and is defined as the time between the beginning of the QRS complex to the end of the T-wave. In Table I below, in which the cardiac cycle is defined as the time between adjacent P waves or QRS complexes. The cardiac cycle is inversely proportional to the heart rate.

The total impedance amplitude determines the simulated stroke volume and is a function of three parameters that change asynchronously during the operation of heart simulator20: a) the pacing interval received from cardiac pacemaker prototyping board30; b) the cardiac cycle time set by the user; and c) the Frank-Sterling factor. The total impedance amplitude is a sum of three contributions:
AmplitudeTotal=AMPPacing—Intervals+AMPCardiac—Cycle+AMPFrank—StarlingEq. 2
AMPPacing—Intervalsdepends on the pacing time interval of each heart beat as defined by AVoptof Eq. 1 above and measured relative to the sensed atrial event. Thus:
PacedAVdelay=AVopt−(Atrial Event(350)−Ventricular Event(370,420))  Eq. 3
AVDelayeffective=PacedAVdelay*1000/Cardiac Cycle  Eq. 4
AMPpacing—Intervals=100−½*|160−AVDelayeffective|  Eq. 5
where the value 100 of Eq. 5 represents the full scale of the AMPPacing—Intervalsamplitude.
AMPCardiac—Cycleis a function of the cardiac cycle and is set forth in Table II below. The variable AMPCardiac—Cyclerepresents the contribution of the heart rate to the simulated to the simulated stroke volume. The stroke volume typically increases with increasing heart rate.

EDZ0 is defined as the value of the impedance sensor when the ventricle is maximally filled, i.e. the end diastolic impedance at the end diastolic volume. EDZ0 is an input parameter selected to shift up or down the full impedance waveform so as to simulate various physiological conditions. As described above the Frank Starling principal defines a relationship between the stroke volume and the end diastolic volume, and indicates that a higher stroke volume is associated with a larger end diastolic volume. Thus adjusting the EDZ0 factor further enables simulating heart activity as a function of heart contractility. The unit31is defined is defined arbitrarily as the maximum value, and the unit50is also defined as the arbitrary maximum value for the Frank Starling amplitude value.

The impedance waveform signal output by hemodynamic sensor model200is composed of 3 or 4 segments according to the states described above in relation toFIG. 4. In each segment the waveform is produced as a straight line for simplicity, however this is not meant to be limiting in any way. The use of curves or a plurality of subsegments within a state is specifically contemplated and is within the scope of the invention. The slope of the line, or lines, within the segment is preferably recalculated every cardiac cycle responsive to user defined parameters and pacing intervals.

The impedance slope during passive filling state300is uniformly negative and is equal to ⅔ of the total amplitude, AmplitudeTotal, divided by the time interval defined by Cardiac Cycle−AVopt−QT Interval, with the variable ⅔selected based on an assumed value that ⅔ of the total diastolic volume is due to passive filling. The impedance slope during active filling state310is negative and equal to ⅓ of the total amplitude, AmplitudeTotal, divided by AVopt. The slope during ejection state320is positive and is equal to total amplitude, AmplitudeTotal, divided by QT Interval. The slope during both regurgitation state330and late passive state340is positive, with the slope being determined as a consequence of the severity of the mitral regurgitation condition being simulated.

The user is preferably provided with an input screen via computing platform40to enter a number of variables. In particular, and without limitation, the user may define the cardiac cycle time, i.e. the atrial rate. The cardiac cycle time defines the time interval between atrial events. The user may also define EDZ0 as described above. The user also may define AVoptat a predetermined heart rate, for example AVoptat 60 beats per minute. Preferably, hemodynamic sensor model200is operative to calculate AVoptas required for the actual heart rate, or cardiac cycle time in the absence of an override input from the user.

The pacemaker inputs to the simulated sensor model are: a) a paced AV interval of the right ventricular measured from the right atrial event; and b) a paced AV interval of the left ventricular measured from the right atrial event. The AV delay is defined as the delay from an atrial event that triggers filling of the atrium to a ventricular event that triggers blood ejection from the ventricle. The VV interval is defined as the interval between the pacing signal for the left ventricle to the pacing signal for the right ventricle. Thus, when the left ventricle is paced before the right ventricle the VV interval is negative. The VV Interval is defined as:
VVinterval=PacedleftAV−PacedrightAV[msec].  Eq. 7

In an exemplary embodiment the hemodynamic sensor model further allows the insertion of a random noise to the output. The random noise is added to the amplitude of Eq. 2 to define an amplitude total with noise, denoted AmplitudeTotalwithnoiseas:
AmplitudeTotalwithnoise=AmplitudeTotal+AMPRandom NoiseEq. 8

The random noise input as AMPRandom Noisemimics all sources of noise that can affect the output of the actual hemodynamic sensors in-vitro. This ability to add random noise advantageously further allows for the testing of cardiac pacemaker prototyping board30, and a control system being tested thereon, in realistic conditions.

FIG. 5illustrates an output of the hemodynamic sensor model simulating an impedance sensor in accordance with the principle of the current invention, the output being indicative of nominal heart operation. The x-axis represents time in milliseconds and the y-axis represents a normalized impedance value. At point500, expiration of a QT timer, which mimics a T-wave, triggers passive filling state300during which the impedance slope is negative from point500to point510. Point510represents sensed atrial event350triggering the transition from passive filling state to active filling state310. The slope from point510is more negative, characteristic of active filling state310which ends at point520, representative of ventricular event370. The delay between point510and point520is the AVdelay. The impedance at point520is at a minimum, indicative of maximum blood volume, and is defined as end diastolic volume (EDV).

The impedance slope from point520is positive, characteristic of ejection state320to point530. Point530is representative of event400, i.e. the expiration of the QT Interval. The impedance at point530is at a maximum indicative of a minimal blood volume, and is defined as end systolic volume (ESV). The stroke volume (SV) is proportional to the impedance differences:
Stroke Volume=f(ZESV−ZEDV)  Eq. 9

FIG. 6aillustrates an output of the hemodynamic sensor model simulating an impedance sensor in accordance with the principle of the current invention, the output being indicative of a too short AV delay. The short AV delay reduces the stroke volume due to sub-optimal too short filling phase. The x-axis represents time in milliseconds and the y-axis represents a normalized impedance value. At point600, a QT timer triggers the passive filling state300during which the impedance slope is negative from point600to point610. Point610represents atrial event350triggering the transition from passive filling state300to active filling state310. The slope from point610is more negative, characteristic of active filling state310which ends at point620, representative of ventricular event370. It is to be noted that point620occurs earlier than desired, i.e. before the optimal AVdelay, AVopt. The impedance at point620is at a minimum, indicative of maximum blood volume, i.e. EDV. Due to early triggering of point620, EDV is not optimal.

The impedance slope from point620is positive, characteristic of ejection state320to point630. Point630is representative of event400, i.e. the expiration of the QT Interval. The impedance at point630is at a maximum indicative of a minimal blood volume, i.e. ESV. It is to be noted that due to the early triggering of point620, SV ofFIG. 6ais lower than the optimal timing SV ofFIG. 5. The output signal as represented byFIG. 6a, and other temporal signals as required, may be further output to computing device40ofFIG. 1to enable the determination of various mechanical hemodynamic information such as, but not limited to, SV.

FIG. 6billustrates an output of the hemodynamic sensor model simulating an impedance sensor in accordance with the principle of the current invention, the output being indicative of a too long AV delay. The long AV delay reduces the SV due to regurgitation. The x-axis represents time in milliseconds and the y-axis represents a normalized impedance value. At point700, a QT timer triggers passive filling state300. The impedance slope is negative from point700to point710characteristic of passive filling state300which ends at point710. Point710represents atrial event350triggering the transition from passive filling state300to active filling state310. The slope from point710is more negative, characteristic of active filling state310which ends at point720, representative of the expiration of AV timer230, i.e. event380which triggers the transition to regurgitation state330. The slope from point720is slightly positive characteristic of regurgitation state330and indicative of some blood escaping the ventrical due to regurgitation. The impedance at point720is at a minimum, indicative of maximum blood volume, i.e. EDV. The impedance increases from point720up to point730indicative of paced ventricular event420, which transitions from regurgitation state330to ejection state320.

The impedance slope from point730is strongly positive, characteristic of ejection state320to point740. Point740is representative of event400, i.e. the expiration of the QT Interval. The impedance at point740is at a maximum indicative of a minimal blood volume, i.e. ESV. It is to be noted that due to the late triggering of point730, the SV is lower than the SV characteristic of the optimal timing ofFIG. 5. The output signal as represented byFIG. 6b, and other temporal signals as required, may be further output to computing device40ofFIG. 1to enable the determination of various mechanical hemodynamic information such as, but not limited to, SV

FIG. 7aillustrates the dependence of the output of the impedance sensors of the hemodynamic sensor model on HR in which the x-axis represents time in milliseconds and the y-axis represents impedance values. Three different impedance waves are shown denoted heart rate1, heart rate2and heart rate3. It is to be noted that heart rate1<heart rate2<heart rate3. With increasing heart rate an increasing stroke volume and impedance range is exhibited.

FIG. 7billustrates the dependence of the output of the impedance sensors of the hemodynamic sensor model on both HR and the end diastolic impedance parameter, EDZ0, in which the x-axis represents time in milliseconds and the y-axis represents impedance values. Three different EDZ0 values are illustrated denoted EDZ01, EDZ02and EDZ03. If is to be noted that EDZ01>EDZ02>EDZ03. Thus as EDZ0 is reduced the entire impedance curve shifts upwards indicative of higher impedance values corresponding to lower ventricular volumes.

Thus the present embodiments enable a heart simulator that generates simulated implanted IEGM signals and hemodynamic sensor signals thereby simulating both heart electrical and mechanical activity, the simulation signals being responsive to the output of a CRT device. Preferably, the heart simulator further exhibits a programming input, allowing for simulation of various heart action irregularities.