Multiple parameter rate-responsive cardiac stimulation apparatus

A metabolic demand rate-responsive cardiac stimulation apparatus and method are disclosed which employ multiple physiological rate control parameters, such as respiratory minute volume, patient motion and cardiac stroke volume. The parameters are derived using a single standard pacing lead or transducer. The apparatus and method perform each physiological measurement by periodically applying a measuring current between two points within the apparatus. This measuring current has frequency components in a range of from approximately 10 kilohertz to 1000 megahertz. Application of this measuring current allows the apparatus to detect the voltage which arises from the applied current and, from the detected voltage, to measure the patient's spatial impedance. For a particular measurement, the apparatus controls which physiological parameter is sensed by regulating the frequency content of the measuring current. The apparatus analyzes the physiological parameters to determine the best pacing rate in terms of characteristics such as response time and stability.

TECHNICAL FIELD 
The present invention relates to metabolic demand rate-responsive cardiac 
control devices, and more particularly to such devices having multiple 
metabolic demand sensors, the outputs of which are analyzed to derive a 
rate-control parameter. 
BACKGROUND OF THE INVENTION 
A general requirement for any rate-responsive pacemaker is a sensor for 
detecting a physiological parameter which varies with the body's metabolic 
demand for cardiac output of blood. Preferably, a rate-responsive 
pacemaker will monitor a physiological parameter which accurately responds 
to physical and emotional stimulation wherein changes in the parameter and 
in metabolic demand vary in a linear fashion. Various types of 
rate-responsive pacemakers have been developed which provide different 
approaches to metabolic-demand sensing. These pacemakers may measure 
different physiological parameters or measure a particular physiological 
parameter in a different manner to provide a basis for rate-adaptive 
pacing. Each of these different approaches to metabolic-demand sensing may 
be advantageous or disadvantageous for a particular patient or cardiac 
malfunction. 
The present invention provides a rate-adaptive sensor within a pacemaker 
which allows the pacemaker to automatically match the pacing rate to the 
patient's metabolic demand and to respond quickly to changes in the 
metabolic demand. The operation of the sensor may be altered by means of 
programming of the pacemaker from an external communicating device. These 
alterations in sensor operation fulfill the needs of various patients who 
are afflicted with different cardiac and respiratory health problems. This 
sensor does not require special leads or special sensor transducers other 
than those common in standard cardiac pacemakers. 
One common metabolic-demand sensor measures physical activity to provide a 
suitable parameter for rate adaptation. A physical activity sensor is not 
generally regarded as a truly physiologic sensor because it does not 
measure true metabolic demand and, therefore, is not affected by emotional 
stimuli or pyrexia. In U.S. Pat. No. 4,140,132, entitled "Variable Rate 
Timer for a Cardiac Pacemaker", issued to J. D. Dahl on Feb. 20, 1979, a 
pacemaker employs an accelerometer, an implanted weighted cantilever arm 
piezoelectric crystal, to monitor the physical activity of a patient and 
set the pacemaker's escape interval. Similarly, in U.S. Pat. No. 
4,428,378, entitled "Rate Adaptive Pacer", issued on Jan. 31, 1984, K. M. 
Anderson et al. describe a sensor which generates a signal reflecting the 
activity of a patient. The pacemaker bandpass filters this signal, detects 
its amplitude and derives a pacing rate from the processed signal. The 
high frequency content of this signal increases with patient movement, 
therefore the pacemaker modulates the pacing rate, between preset rate 
maxima and minima, in proportion to the intensity of the processed signal. 
The most important advantage of a physical-activity sensor is its very 
rapid response time to the onset of exercise. A physical-activity sensor 
responds favorably to patient activities which create vibration, such as 
jogging, walking and stair climbing. Unfortunately, activities such as 
bicycling do not promote rate adaptation because little vibration occurs. 
Further advantages of a physical-activity sensor lie in its simplicity. No 
special pacing lead is required since an activity-based pacemaker may 
employ standard leads with either unipolar or bipolar electrodes. 
Furthermore, no special implanting procedure is required for an 
activity-based pacemaker. 
Although the lack of a truly physiological response is generally considered 
a disadvantage of an activity sensor, the fact that this sensor acts 
independently from physiologic variables may provide a better response 
under conditions in which patient systems or tissues are diseased. For 
example, an activity sensor may supply a better signal for responding to 
exercise than a respiration sensor will under conditions of lung disease, 
such as emphysema. 
The primary disadvantage of a rate-responsive pacemaker employing a 
physical-activity sensor is the difficulty of attaining a scaled response 
to gradations of metabolic demand. Activity sensors generally act in an 
on/off fashion, in which a sensor is unable to detect changes in patient 
workload. Therefore, the response of activity-based, rate-responsive 
pacemakers does not normally depend on the amount of exercise the patient 
is performing, but instead the rate change remains identical so long as 
the measured activity is above a preprogrammed level. Because it is 
difficult or impossible to relate the amount of vibration of the sensor to 
the cardiac output needs of a patient performing activity of various 
types, the sensor cannot be programmed to adapt the pacing rate in a 
physiological manner. In particular, it is difficult to program the 
pacemaker to correctly respond to the onset or cessation of exercise. 
Furthermore, a physical-activity sensor generates an undesirable response 
to noise disturbances arising external to the body (e.g., machinery) or 
from within the body (e.g., coughing, sneezing and laughing). Also, noise 
signals tend to swamp activity-induced signals which occur at some 
frequencies. 
A second type of metabolic-demand sensor measures and analyzes impedance 
signals which relate to cardiac mechanical performance to adapt the pacing 
rate to the metabolic demands of the patient. Pacemakers analyze and 
process cardiac mechanical data to derive physiological parameters such as 
stroke volume or cardiac output. For example, a pacemaker may utilize an 
intravascular-impedance sensor to measure right ventricular stroke volume 
and adjust pacing rate to keep this parameter at predetermined physiologic 
values. In U.S. Pat. No. 4,535,774, entitled "Stroke Volume Controlled 
Pacer", issued on Aug. 20, 1985, W. H. Olson describes an impedance 
plethysmography sensor, comprising a number of electrodes and analysis 
circuitry, which is employed to detect variations in stroke volume over 
time. The pacemaker sets the pacing rate according to these changes in 
stroke volume. 
The primary advantage of using stroke volume as a parameter for adjusting 
rate in a rate-responsive pacemaker is the capability of rapidly adjusting 
the rate to changes in metabolic demand in a physiologic manner. 
The main disadvantage of the stroke-volume, rate-responsive pacemaker is 
its requirement for a nonstandard pacing lead having multiple electrodes. 
Sensing of impedance using standard bipolar leads has not provided the 
accuracy in the stroke volume measurement which is necessary for 
rate-adaptive control. Preferable tripolar or quadripolar leads have not 
been durable enough for chronically implanted usage. Furthermore, an 
appropriate algorithm for driving a closed loop adaptive pacing rate has 
not been discovered. Rate-response algorithms using stroke volume as a 
control parameter have been most ineffective for sick patients. A 
disadvantage of stroke-volume controlled rate adaptation, in comparison to 
rate control based on physical activity, is its requirement for complex 
sensors and circuits and an inability to use standard pacing leads. 
Another type of metabolic-demand sensor measures and analyzes impedance 
signals which relate to a patient's respiratory function to adjust pacing 
according to the metabolic demands of the patient. 
In U.S. Pat. No. 4,567,892, entitled "Implantable Cardiac Pacemaker", 
issued to G. Plicchi and G. Canducci on Feb. 4, 1986, a pacemaker is 
disclosed which monitors respiratory rate by measuring impedance 
variations throughout a distance within the thoracic region of a patient's 
body, between the pacemaker can and a separate auxiliary or passive lead 
implanted subcutaneously in the chest wall using a special tunneler. A 
programmed algorithm within the pacemaker analyzes the respiratory rate to 
determine a pacing rate. It has been shown that heart rate, respiratory 
rate and oxygen uptake all correlate well irrespective of the presence of 
lung disease. All parameters increase at the onset of exercise and 
decrease when exercise stops. The rate-responsive pacemaker which is 
driven by the respiration rate measurement is simple and reliable, as well 
as sound in its physiologic basis. 
A fundamental disadvantage of driving the pacing rate on the basis of 
respiratory-rate variations is that it takes into account only part of the 
body's ventilation adaptation in response to exercise. Ventilation 
increases due to an increase in the depth of respiration as well as the 
respiration rate. Although respiratory rate relates somewhat closely to 
heart rate, heart rate correlates much more strongly with the total amount 
of inspired air. That the Plicchi and Canducci pacemaker requires a 
special surgical procedure, tunneling of a lead in the patient's thoracic 
region, and a special sensor, are practical disadvantages of the device. 
M. S. Lampadius, in U.S. Pat. No. 4,721,110, entitled 
"Respiration-controlled Cardiac Pacemaker", issued on Jan. 26, 1988, 
improves on the respiration-rate-driven pacemaker by disclosing a 
pacemaker driven either by respiration depth or respiration rate. This 
pacemaker employs a rheography pulse generator, which generates constant 
amplitude pulses during the refractory period of a patient's heart, and a 
respiration detector which, as a function of the impedance data measured 
in response to the rheography pulses, generates a respiration signal 
representing the respiratory rate, the depth of respiration or a 
combination of the rate and depth of respiration. The pacemaker then uses 
this respiration signal to determine an appropriate pacing stimulation 
rate. 
The respiratory parameter which correlates most closely to heart rate is 
minute ventilation, a highly physiologic variable which reflects closely 
the metabolic demands of exercise. The body's increase in minute 
ventilation during exercise parallels its oxygen uptake but also reflects 
changes in cardiac output and heart rate. Minute ventilation not only 
correlates well with exercise, but also varies in response to stress and 
pyrexia. U.S. Pat. No. 4,702,253 (hereinafter called the "'253 patent"), 
entitled "Metabolic-Demand Pacemaker and Method of Using the Same to 
Determine Minute Volume", issued to T. A. Nappholz et al. on Oct. 27, 
1987, discloses a rate-responsive pacemaker which senses impedance in the 
pleural Cavity of a patient and derives respiratory minute volume from 
impedance. The pacemaker then employs the respiratory minute volume, a 
measure of the amount of air inspired by a person as a function of time, 
as a rate-control parameter. The greater the amount of air inspired, the 
greater the need for a higher pacing rate. The device described in this 
patent requires a nonstandard pacing lead in order to perform the minute 
volume measurement. 
U.S. Pat. No. 4,901,725 (hereinafter called the "'725 patent"), entitled 
"Minute Volume Rate-Responsive Pacemaker", issued to T. A. Nappholz et al. 
on Feb. 20, 1990, discloses a pacemaker which performs a rate-responsive 
function in the manner of the '253 patent with various improvements, and, 
in addition, only requires standard pacing leads. To measure the 
intravascular impedance, the minute-volume sensor generates a low energy 
current pulse at 50 ms intervals between a ring electrode of the lead and 
the pulse generator case, then measures the voltage between the tip 
electrode of the lead and the pulse generator case arising from the 
applied current. An intravascular impedance value is determined from the 
measured voltage and the applied current using Ohm's law. Transthoracic 
impedance increases with inspiration, decreases with expiration and its 
amplitude varies with the tidal volume. The impedance signal thus 
comprises two components, representing tidal volume and respiratory rate. 
Pulse generator circuitry identifies the two signals and processes them to 
yield minute ventilation. The minute-volume-controlled rate-responsive 
pacemaker employs a highly physiologic sensor. Its ability to assess the 
metabolic demands of the body are superior to that of a pacemaker driven 
by respiratory rate alone, since depth of ventilation is an important 
response to exercise or stress. The apparatus described in the '725 patent 
requires no more than standard pacing leads, although the leads cannot be 
unipolar leads, and programming of minute ventilation rate adaptation 
necessitates only a single exercise test. 
Pacemakers which use any of the discussed respiratory parameters as a basis 
for rate adaptation are considered to respond more slowly to the onset of 
exercise than a physical activity controlled pacemaker. A faster response 
is more desirable. 
An advantage of the activity-sensing pacemakers is its fast response to the 
onset of exercise, but a major disadvantage of a pacemaker which 
determines pacing rate based on an activity signal alone is its 
substantial inability to react to the instantaneous metabolic level of 
exercise or stress. A pacemaker which uses a more physiologic parameter 
may respond less quickly to the onset of exercise, but is highly specific 
with respect to the metabolic level of exercise. Thus, a pacemaker may 
employ two parameters in combination in a rate adaptive cardiac pacing 
system, wherein each parameter complements the other by mutually supplying 
what the other lacks. 
U.S. Pat. No. 4,926,863, entitled "Rate-responsive Cardiac Pacemaker", 
issued to E. Alt on May 22, 1990, discloses a rate-responsive cardiac 
pacemaker which employs an activity sensor, in the form of an 
accelerometer, as a first sensor of metabolic demand and a second sensor 
which is adapted to detect a parameter which is complementary to activity. 
The measurement of the second sensor is used to supply a "complementary 
parameter" for confirming the presence of a particular metabolic state and 
selectively contributing to the determination of a stimulation rate. The 
complementary parameter of the second sensor is defined as any 
physiological or other detected parameter, within or outside the body, 
having characteristics of sensitivity and specificity to physical exercise 
which contrast and enhance the corresponding characteristics of the 
activity sensor, specifically, a fast response time to the onset of 
exercise but nonspecificity with respect to the instantaneous metabolic 
level of exercise. The Alt patent mentions the parameter of central venous 
blood temperature as a possible complementary parameter. 
U.S. Pat. No. 4,860,751, entitled "Activity Sensor for Pacemaker Control", 
issued to F. J. Callaghan on Aug. 29, 1989, also discloses a 
rate-responsive cardiac pacemaker which includes an activity sensor in 
combination with a second sensor for monitoring a physiological parameter 
such as partial pressure of oxygen (pO.sub.2), blood pressure, core 
temperature, CO.sub.2 and pCO.sub.2, O.sub.2 and pO.sub.2, pH, respiration 
rate, respiration depth and ventricular volume. In this patent, the 
activity sensor provides for generally constant monitoring of patient 
motion. When patient motion causes the activity sensor to generate a 
signal exceeding a preset threshold level, the physiological sensor is 
activated. The parameter generated by the physiological sensor is used to 
set pacing rate. The advantage of using an activity sensor as a trigger to 
initiate sensing by a physiological sensor is that the activity sensor 
requires very little energy expenditure to operate. In contrast, a 
physiological parameter sensor normally consumes a great deal of energy. 
SUMMARY OF THE INVENTION 
In accordance with one aspect of the present invention, there is provided a 
metabolic-demand, rate-responsive cardiac stimulation apparatus which 
paces a patient's heart at a controlled rate. The apparatus analyzes a 
signal from an impedance-measuring sensor within the patient's body to 
determine this controlled rate. This sensor applies multiple fixed 
frequencies of current between two points within the apparatus and 
measures voltages in response to the application of the measuring 
currents. The measuring current is in the form of a measuring signal 
having frequency components within a range from approximately 10 kilohertz 
to approximately 1000 megahertz. The apparatus includes a means for 
controlling the impedance sensor, which limits the frequency components of 
the applied time-varying current to lie within at least one predetermined 
subrange of frequencies. For each subrange of measuring frequencies, the 
measured voltage arising in response to the applied current corresponds to 
a measurement of a particular metabolic-demand parameter. The apparatus 
further includes a means to derive at least one metabolic-demand parameter 
from the measured voltage corresponding to each predetermined subrange of 
frequencies. A means for determining the controlled pacing rate sets the 
rate in terms of a predefined relationship or rate to the values of at 
least one of the derived metabolic-demand parameters. 
The means for pacing a patient's heart at a controlled rate within the 
apparatus of the present invention includes at least one pacing lead. In 
some embodiments of the present invention, the impedance-measuring sensor 
applies a measuring current and measures the voltage resulting from the 
applied measuring current by means of at least one electrical coupling to 
a pacing lead. The other electrical coupling of the impedance-measuring 
sensor may either connect with a point within the apparatus, such as its 
case, or may connect with another lead, if one is included within the 
apparatus. 
The apparatus may also include a transmission coil, for example, a coil 
that is commonly used to perform telemetric communication. In embodiments 
of the invention which are an alternative to embodiments in which the 
impedance-measuring sensor couples to at least one lead, this sensor may 
couple to the coil to generate measuring current and measure the resulting 
voltage. An implementation of the apparatus which employs a coil to sense 
impedance provides a different signal for analysis than an implementation 
employing a pacing lead. The coil can measure mechanical movement of the 
heart alone, whereas the lead measures mechanical movement in combination 
with blood flow. (Note that a lead which is insulated at its tip also 
senses mechanical movement of the heart without a signal component 
relating to blood flow.) A coil which is affixed within the pacemaker case 
can still sense impedance. 
A coil may be affixed to the end of a pacemaker lead to sense impedance in 
its vicinity. The signal produced by sensing impedance from a coil at the 
end of a lead differs from the signal produced by sensing impedance from 
the lead itself, since the lead will sense impedance along its entire 
length while the coil will sense only locally, at the tip of the lead. 
The impedance-measuring sensor generates measuring current in the form of 
either short-duration, square-wave-like current pulses, sinusoidal-like 
oscillating current or short-duration pulses of sinusoidal-like 
oscillating current. Current pulses range in duration from 5 ns to 20 
.mu.s. Oscillating current ranges in frequency from 10 kHz to 1000 MHz, 
and may be generated continuously or in pulses as short as 5 ns. 
If the impedance-measuring sensor generates measuring current on a pacing 
lead in the form of short-duration, square-wave-like current pulses or 
short-duration pulses of sinusoidal-like oscillating current, the sensor 
control means sets the value of the duration for the purpose of measuring 
a particular metabolic-demand parameter. One possible metabolic-demand 
parameter is based on patient motion. Its associated subrange of durations 
includes pulse durations shorter than approximately 125 nanoseconds. A 
second possible metabolic-demand parameter is based on respiration and its 
associated subrange of durations includes pulse durations from 
approximately 50 nanoseconds to approximately 400 nanoseconds. A third 
possible metabolic-demand parameter is based on cardiac hemodynamic 
signals. Its associated subrange of durations includes pulse durations 
longer than approximately 300 nanoseconds. 
If the impedance measuring sensor generates measuring current on a pacing 
lead in the form of sinusoidal-like oscillating current, the sensor 
control means sets the value of the sinusoidal-like oscillating frequency 
for the purpose of measuring a particular metabolic-demand parameter. One 
possible metabolic-demand parameter is based on patient motion. Its 
associated subrange of frequencies includes those higher than 
approximately 8 megahertz. A second possible metabolic-demand parameter is 
based on respiration and its associated subrange of frequencies ranges 
from approximately 1 megahertz to approximately 11 megahertz. A third 
possible metabolic-demand parameter is based on cardiac hemodynamic 
signals. Its associated subrange of frequencies is the range less than 
about 4 megahertz. 
In some embodiments of the invention, if the impedance measuring sensor 
generates measuring current in the form of pulses, the sensor control 
means sets the value of the duration for the purpose of measuring a 
particular metabolic-demand parameter. The sensor control means may 
consistently set the duration to a single value to measure only one 
metabolic parameter. Alternatively, the sensor control means may set a 
number of duration values and interleave (in time) current pulses with 
different durations to measure multiple metabolic parameters. The sensor 
control means may time the different duration pulses to take place in any 
sequence so that different sampling rates may apply to the pulse of 
varying duration. 
In embodiments of the invention which are alternative to embodiments in 
which the sensor control means varies the measuring-pulse duration, the 
sensor control means generates short-duration, square-wave-like current 
pulses or short-duration pulses of sinusoidal-like oscillating current to 
a particular predefined duration and sets the timing of each voltage 
measurement for the purpose of measuring a particular metabolic-demand 
parameter. For each current pulse, the impedance-measuring sensor may 
sample the voltage at one or more times. These sampling times are set by 
the sensor control means. Each of these voltage sampling times defines a 
sampling duration which begins at the onset of the current pulse. The 
sensor control means may set multiple sampling durations to allow the 
impedance-measuring sensor to measure multiple metabolic parameters. The 
sampling durations in these alternative embodiments of the invention 
produce signals which are equivalent to the same measuring-current-pulse 
durations in embodiments of the invention in which the impedance-measuring 
sensor generates measuring-current pulses of different durations. 
The present invention incorporates two or more rate-adaptive sensors in a 
single rate-responsive pacing apparatus. These multiple sensors reside in 
a single circuit which requires only standard bipolar and unipolar pacing 
leads, rather than special, nonstandard pacing leads. The multiple sensors 
do not need special sensor transducers other than the sensing circuits 
which are common in the field of cardiac pacemakers. Appropriate selection 
of these sensors provides for automatic matching of the pacing rate to the 
patient's metabolic demand and to respond quickly to changes in the 
metabolic demand. The selection and control of these sensors may be 
altered by means of programming of the pacemaker from an external 
communicating device. Thus, a physician may activate one or more sensors 
to provide a rate-adaptive parameter which is best suited to a patient, 
thereby fulfilling the needs of various patients who are afflicted with 
different cardiac and respiratory health problems. 
The present invention provides a single pacing apparatus in which a single 
impedance-measuring sensor may measure completely separate 
metabolic-demand parameters employing independent analysis procedures. A 
physician may select the metabolic parameter and analysis method most 
suitable for a patient. Alternatively, the physician may enable a more 
complex analysis system in which the sensed metabolic-demand parameters 
are combined via a control logic such that deficiencies in the response to 
one sensed parameter can be overcome by another. For example, rapid 
response of activity may be combined with stable response of minute 
ventilation.

DETAILED DESCRIPTION 
The drawing of FIG. 1 is a high-level block schematic of the apparatus of 
the invention in the form of a pacemaker, shown generally at 5. All 
pacemaker logic is under the control of a controller 28 (which may include 
a microprocessor, although discrete blocks are shown in FIG. 14). The 
controller operates various switches in the pacemaker, of which only one 
pair SW1a, SW1b is shown. Switch SW1b is closed whenever the pacemaker is 
to pace or sense. 
Referring to FIGS. 1 and 2, together, in order to pace, the controller 28 
sends a command to a pulse generator 18 by means of a signal on a E 
conductor 26. The pulse generator 18 responds to this command by applying 
a current pulse through the switch SW1b and a conventional unipolar pacing 
lead 11 to the latter's tip electrode 10, which is shown positioned in the 
right ventricle 12 of a patient's heart 7 in FIG. 2. A sense amplifier 16 
senses a cardiac signal on the electrode. (Various functions well known in 
the art, such as blanking of the sense amplifier during pacing, are not 
shown inasmuch as they have no bearing on the subject invention.) The 
sensing of a heartbeat, spontaneous or evoked, results in a pulse 
appearing on a SENSE conductor 24 for communication to controller 28. The 
"SENSE" function activates the loading of an initial "count" value 62 
(FIG. 18), as will be described hereinafter. 
The pacemaker 5 makes an impedance measurement when the controller 28 
pulses the conductor 20 to activate impedance measurement circuit block 
14. Upon this event, switch SW1a closes, switch SW1b opens and impedance 
measurement circuit block 14 applies a current to the lead 11, causing 
current to flow through the lead toward the tip electrode 10. The 
measuring current which is applied to the electrode has frequency 
characteristics in the range from about 10 kHz to about 1000 MHz. At these 
measuring current frequencies, the lead acts as an antenna which creates a 
displacement current in the body. This displacement current is 
fundamentally different from the conduction current which is generated by 
prior art impedance-measuring pacemakers such as those shown in the 
aforementioned '253 and '725 patents. The impedance measurement circuit 
block 14 may generate this measuring current in the form of 
sinusoidal-like oscillating current, short-duration pulses of 
square-wave-like current, or timed pulses of sinusoidal-like oscillating 
current. The impedance measurement circuit block 14 measures spatial 
impedance by determining the potential between the pacemaker case 30 and 
the pacemaker's input connection to the conductor (not shown) within lead 
11. This conductor extends to the tip electrode 10. In this configuration, 
the pacemaker case 30 serves as a reference potential for the pacemaker 
circuitry. In the preferred embodiment of the pacemaker, the impedance 
measurement circuit block 14 derives samples at a rate of about 20 per 
second and communicates these samples to controller 28 over conductor 22. 
The impedance measurement can be executed as described in the 
aforementioned '253 patent. 
Placement of the unipolar lead 11 is shown in FIG. 2. The tip electrode 10 
makes contact with the wall of the right ventricle 12 or the right atrium 
13 of the patient's heart 7. When the impedance measurement circuit block 
14 generates measuring currents at appropriate frequencies, as will be 
described hereinafter, the impedance measurement reflects minute volume to 
a much greater extent than stroke volume or motion artifacts. In addition, 
the impedance measurement reflects minute volume more than signals 
originating from other physiological and non-physiological sources because 
of the characteristics of a filter 23 (FIG. 3) which is part of the 
impedance measurement circuit (see, also, FIG. 1 of the '253 patent). In 
an embodiment of the invention which filters the impedance signal to favor 
sensing of a respiration signal component over other components, the 
impedance signal is filtered by a two-pole filter with a center frequency 
of 0.2 Hz. The gain is reduced by a factor of two (6 dB) at frequencies of 
0.05 Hz and 0.8 Hz. Alternatively, the capacitors and resistors of the 
circuit of filter 23 may be chosen to preferentially elicit other 
physiological signal components. For example, the cutoff frequencies for a 
bandpass filter which favors cardiac hemodynamic signals may range from 
0.2 to 10 Hz. Furthermore, the cutoff frequencies for a bandpass filter 
which best supports patient motion signals may range from 5 to 10 Hz. 
Referring to FIG. 3, an embodiment of the impedance measurement circuit 14 
which operates in a pulsed mode is shown. The impedance measurement 
circuit 14 includes a connection through a switch SW4 with the pacemaker 
case 30, and a connection through a switch SW3 with the tip electrode 10 
(via the pacing lead 11). The tip electrode 10 is a conventional 
pacing/sensing electrode. The indifferent electrode is the case 30. The 
impedance measurement circuit 14 employs the tip electrode 10 and lead 11 
both for applying a source measuring current to the patient's body, and 
for measuring the impedance between the tip electrode 10, lead 11 and the 
case 30. A buffer 32 (which is discussed hereinafter) and the filter 23 
are also employed in circuit 14. 
All switches in FIG. 3 are directly or indirectly under the control of 
controller 28. One output 33 of the controller is shown extended to switch 
SW2, but it is to be understood that the switches SW3, SW4, SW5, SW6, SW1a 
and SW1b are similarly controlled. The controller closes switch SW6 to 
charge a measuring capacitor C2 to a regulated voltage source VDD. 
Subsequently, the controller opens switch SW6 and closes switches SW3 and 
SW4, while holding switch SW5 open, for a predetermined measuring interval 
.DELTA.T, thereby connecting capacitor C2 to lead 11 through a coupling 
capacitor C3. While the switches SW3 and SW4 are closed, measuring 
capacitor C2 discharges through capacitor C3 into the lead 11, thereby 
decreasing the voltage across measuring capacitor C2. The amount by which 
the voltage across the measuring capacitor C2 diminishes depends on the 
impedance of the lead-tip combination and the impedance of the surrounding 
tissue. The impedance of the lead-tip combination is known and the 
impedance of the surrounding tissue is the object of the measurement. 
Measuring capacitor C2 stores the diminished voltage and buffer 32 later 
transfers this to the measuring circuit in the following manner. After the 
predetermined measuring time interval .DELTA.T, the controller 28 opens 
switches SW3 and SW4, allowing the buffer 32 to access the voltage held on 
the measuring capacitor C2. This voltage is advanced through the buffer 
amplifier 32 and switch SW2 (which the controller 28 closes at the time it 
opens switches SW3 and SW4), and is sampled on capacitor C1 at the input 
of the filter 23. The controller 28 holds switch SW2 closed for a time 
duration which is sufficient for a delta modulator 25 to convert the 
signal into a digital form. For example, a delta modulator which is 
capable of low current operation in an implantable device may commonly 
convert a signal to digital form in about 1 millisecond. 
Switch SW1a is closed during the previously described impedance 
measurements. However, the controller 28 may occasionally command the 
performance of a noise measurement by opening switches SW1b, SW5, SW6 and 
while closing switches SW1a, SW4, and SW2. The controller 28 may thus 
acquire a noise measurement and compare the noise signal with an impedance 
measurement signal to evaluate errors resulting from noise. 
Resuming consideration of the impedance measurements, after converting the 
sample to digital form, the controller 28 opens switch SW2 and closes 
switch SW6 to charge measuring capacitor C2 for the next measurement 
cycle. In the preferred embodiment of the invention, the controller 28 
measures impedance twenty times per second. For each measurement, the 
controller closes the switches SW3 and SW4 for a pulse duration of 250 ns, 
during which the voltage across the measuring capacitor C2 is placed on 
the lead 11. The resistors and capacitors associated with filter 23 pass 
frequencies between about 0.05 Hz and 0.8 Hz, the standard range for 
respiratory measurements. 
The value of the measuring capacitor C2 is selected to store the range of 
voltages which result from various body impedances. In one embodiment of 
the invention, C2 has a capacitance of 4.7 nF. 
The coupling capacitor C3 provides DC isolation for the input to the 
measuring circuit. In one embodiment of the invention, a coupling 
capacitor C3 has a value of about 7.5 .mu.F, which effectively eliminates 
the influence of the DC voltage on measurement results. 
As indicated earlier, the analog signal output of filter 23 passes to delta 
modulator 25 which provides a digital signal output on conductor 22. The 
digital signal output on conductor 22 is input to controller 28 for 
processing, as is hereinafter described in connection with a discussion of 
FIG. 14. Converting an analog signal to a digital representation by delta 
modulation is a standard technique. One example of such an operation is 
illustrated in U.S. Pat. No. 4,692,719 to Robert H. Whigham, entitled 
"Combined Pacemaker Delta Modulator and Bandpass Filter", which issued on 
Sep. 8, 1987. The output of delta modulator 25 is a summation of a series 
of O's and 1's which reflect whether the analog signal is decreasing or 
increasing. 
During a measurement interval, controller 28 opens switch SW1b (shown in 
FIG. 1) to briefly disable pace and sense functions. Although sensing is 
disabled while the impedance measurement is in operation, the duration of 
the measurement is on the order of fractions of microseconds, a time so 
short relative to that of heart signals that disabling sensing during this 
time is of no importance. 
Referring now to FIG. 4 waveform timing diagrams are shown which indicate 
one method by which the controller 28 may regulate switches SW2-SW4 and 
SW6 within the impedance measurement circuit 14 to measure three different 
physiological parameters--patient motion (also called activity), 
respiration and cardiac signals indicative of stroke volume. Referring to 
line (a) of FIG. 4, controller 28 samples respiration, patient motion and 
stroke volume by means of sampling pulses 74, 75 and 76, respectively. The 
controller regulates continuous, interleaved, sampling of each of the 
three physiological parameters, sequentially, at 50 millisecond intervals 
for each type, and with 16.67 milliseconds separating the beginning of 
each sample from adjacent dissimilar samples. The selection of a 50 
millisecond sampling interval is made to illustrate the operation of the 
switches of FIG. 3 and the method of sampling. It is to be understood that 
other sampling intervals are intended to be included within the scope of 
the invention. For example, the various physiological signal components 
are likely to be sampled at different intervals. Cardiac hemodynamic 
signals may be sampled at a higher rate (e.g. 10 millisecond intervals) 
but only during the time blood is ejected from the heart following 
depolarization of the heart (e.g., from 80 milliseconds to 280 
milliseconds after an R-wave). 
Controller 28 has complete control of the sampling procedure. For example, 
the controller may enable or disable the sampling for any of the 
parameters, may individually change the intervals between samples relating 
to a particular parameter or may change the ratio of sampling for one 
parameter in relation to another. 
Lines (b1)-(b3) of FIG. 4 illustrate one example of a procedure by which 
controller 28 may regulate the sampling of the respiration parameter. The 
controller starts the sampling procedure at t.sub.1 by opening switch SW6 
((line (b.sub.3) goes low)) and closing switches SW3 and SW4 ((line 
(b.sub.2) goes high)) for 250 nanoseconds. Next, at t.sub.1 +250, the 
controller 28 opens switches SW3 and SW4 ((line (b2) goes low)) and closes 
switch SW2 ((line (b1) goes high)) to allow delta modulator 25 (of FIG. 3) 
to convert the signal to a digital number. Switch SW2 is held closed for 
the time duration required to digitize the sample (for example, 1 
millisecond). The controller then opens switch SW2 and closes switch SW6 
(action not shown) to finish the sampling procedure and charge the 
capacitor C2 (of FIG. 3). 
In the same manner, lines (c1)-(c3) of FIG. 4 depict an example of the 
procedure by which controller 28 regulates the sampling of the patient 
motion parameter. The controller starts the sampling procedure at t.sub.2 
by opening switch SW6 ((line (c3) goes low)) and closing switches SW3 and 
SW4 ((line (c2) goes high)) for 25 nanoseconds. At the end of the 25 
nanoseconds, switches SW3 and SW4 open ((line (c2) goes low)) and switch 
SW2 closes ((line (c1) goes high)). 
In the same manner, lines (d1)-(d3) of FIG. 4 depict an example of the 
method by which the controller 28 regulates the sampling of the stroke 
volume parameter. The controller 28 starts the sampling procedure at 
t.sub.3 by opening switch SW6 ((line (d3) goes low)) and closing switches 
SW3 and SW4 ((line (d2) goes high)) for 500 nanoseconds. At the end of the 
500 nanoseconds, switches SW3 and SW4 open ((line (d2) goes low)) and 
switch SW2 closes ((line (d1) goes high)). 
Referring now to FIG. 5, waveform timing diagrams are shown which indicate 
a second method by which controller 28 may regulate the switches SW2-SW4 
and SW6 within the impedance measurement circuit 14 to measure the three 
different physiological parameters discussed with respect to FIG. 4. The 
method of FIG. 5 requires a much faster delta modulator 25 than the method 
of FIG. 4. A delta modulator which is capable of performing the FIG. 5 
method must digitize a signal in 200 nanoseconds or less. Referring to 
line (a) of FIG. 5, the controller samples respiration, patient motion and 
stroke volume by generating a current pulse 70, beginning at leading edge 
70a and lasting a duration of 500 nanoseconds. The controller then directs 
the measurement of patient motion, respiration and stroke volume by 
sampling the resulting voltage at the times shown in FIG. 5 by lines 71, 
72 and 73, respectively. The controller 28 may regulate continuous 
sampling of each of the three physiological parameters, sequentially, at 
preselected intervals (possibly 50 millisecond intervals, as was done in 
FIG. 4). In a manner similar to the description given with respect to FIG. 
4, controller 28 has complete control of the sampling procedure. 
Lines (b1)-(b3) of FIG. 5 illustrate one example of a procedure by which 
controller 28 regulates the sampling of the patient motion, respiration 
and stroke volume parameters. The controller starts the sampling procedure 
by opening switch SW6 ((line (b3) goes low)), as shown at 70b, and closing 
switches SW3 and SW4 ((line (b2) goes high)), as shown at 70c, for a 
duration which is long enough to sample any of the desired parameters (for 
example, 500 nanoseconds to measure stroke volume 73). The controller 
times the duration of the shortest sampling measurement duration, for 
example 25 nanoseconds to sample the patient motion parameter 71. Next, 
the controller closes switch SW2, as shown at 71a, to allow the delta 
modulator 25 (of FIG. 3) to convert the signal to a digital number. SW2 is 
held closed for the time duration required to digitize the sample (for 
example, 200 nanoseconds) and then opens, as shown at 71b. In this 
example, the controller 28 holds switch SW2 open for 25 nanoseconds, as 
shown between 71b and 72a, after which time the measuring current pulse 
has been applied for 250 nanoseconds, the time duration of the respiration 
measurement 72. Again, the controller 28 closes switch SW2, as shown at 
72a, for 200 nanoseconds to allow the delta modulator 25 (of FIG. 3) to 
convert the signal to digital form. The controller 28 then opens switch 
SW2, as shown at 72b, for 50 nanoseconds, after which time the measuring 
current pulse has been applied for 500 nanoseconds, the time duration of 
the stroke volume measurement 73. Next, the controller 28 opens switches 
SW3 and SW4 and closes switch SW2, as shown at 73a, and the delta 
modulator 25 (of FIG. 3) converts the stroke volume signal to a digital 
number. Again, the controller 28 holds SW2 closed 200 nanoseconds. The 
controller 28 then opens switch SW2, as shown at 73b, and closes switch 
SW6, as shown at 73c, to finish the sampling procedure and charge the 
capacitor C2 (of FIG. 3). 
From the foregoing description, it is apparent that the circuit of FIG. 3 
is a component of an embodiment of the present invention in which a single 
sensor and circuit is capable of measuring multiple physiological 
parameters. The timing diagrams of FIG. 5 illustrate that the voltage 
arising from a single measuring current pulse may be sampled at various 
times to measure distinct and separate physiological parameters. 
Referring to FIG. 6, an embodiment of the impedance measurement circuit 14, 
which operates in a manner similar to the circuit of FIG. 3, is shown. The 
impedance measurement circuit 14 includes a connection through a switch 
SW4 with the case 30, and connections through switches SW3, SW13 and SW23 
with the tip electrode 10 (via the pacing lead 11). As in the case of FIG. 
3, the tip electrode 10 is a conventional pacing/sensing electrode and the 
indifferent electrode is the case 30. Buffers 32, 41 and 43 and filters 
23, 34 and 39 (which are discussed hereinafter) are also employed in 
circuit 14. 
All switches in FIG. 6 are controlled by controller 28. One output 33 of 
the controller is shown extended to switch SW2, but it is to be understood 
that the switches SW3, SW4, SW5, SW6, SW12, SW13, SW16, SW22, SW23 and 
SW26 are similarly controlled. The controller closes switches SW6, SW16 
and SW26 to charge measuring capacitors C2, C4 and C5 to a regulated 
voltage source VDD. Subsequently, the controller opens switches SW6, SW16 
and SW26 and closes switches SW3 and SW4, while holding switch SW5 open, 
for a predetermined measuring interval .DELTA.T.sub.1, thereby connecting 
capacitor C2 to lead 11 through a coupling capacitor C3. While the 
switches SW3 and SW4 are closed, measuring capacitor C2 discharges through 
capacitor C3 into the lead 11, thereby decreasing the voltage across 
measuring capacitor C2. The amount by which the voltage across the 
measuring capacitor C2 diminishes depends on the impedance of the lead-tip 
combination and the impedance of the surrounding tissue. The impedance of 
the lead-tip combination is known and the impedance of the surrounding 
tissue is the object of the measurement. Measuring capacitor C2 stores the 
diminished voltage and buffer 32 later transfers this to the measuring 
circuit in the following manner. After the predetermined measuring time 
interval .DELTA.T.sub.1, the controller 28 opens switches SW3 and SW4, 
allowing the buffer 32 to access the voltage held on the measuring 
capacitor C2. This voltage is advanced through the buffer amplifier 32 and 
switch SW2 (which the controller 28 closes at the time it opens switches 
SW3 and SW4), and is sampled on capacitor C1 at the input of the filter 
23. The controller 28 holds switch SW2 closed for a time duration which is 
sufficient for delta modulator 25 to convert the signal into a digital 
form. After converting the sample to digital form, the controller 28 opens 
switch SW2 and closes switch SW6 to charge measuring capacitor C2 for the 
next measurement cycle. 
Next, the controller 28 closes switches SW13 and SW4, while holding switch 
SW5 open, for a second predetermined measuring interval .DELTA.T.sub.2, 
thereby connecting capacitor C4 to lead 11 through the coupling capacitor 
C3. While the switches SW13 and SW4 are closed, measuring capacitor C4 
discharges through capacitor C3 into the lead 11. Measuring capacitor C4 
stores the voltage which buffer 41 later transfers to the measuring 
circuit. After the predetermined measuring time interval .DELTA.T.sub.2, 
the controller 28 opens switches SW13 and SW4, allowing the buffer 41 to 
access the voltage held on the measuring capacitor C12. This voltage is 
advanced through the buffer amplifier 41 and switch SW12 (which the 
controller 28 closes at the time it opens switches SW13 and SW4), and is 
sampled on capacitor C6 at the input of the filter 34. The controller 28 
holds switch SW12 closed for a time duration which is sufficient for delta 
modulator 25 to convert the signal to digital form. After converting the 
sample to digital form, the controller 28 opens switch SW12 and closes 
switch SW16 to charge measuring capacitor C4 for the next measurement 
cycle. 
Next, the controller 28 closes switches SW23 and SW4, while holding switch 
SW5 open, for a third predetermined measuring interval .DELTA.T.sub.3, 
thereby connecting capacitor C5 to lead 11 through the coupling capacitor 
C3. While the switches SW23 and SW4 are closed, measuring capacitor C5 
discharges through capacitor C3 into the lead 11. Measuring capacitor C5 
stores the voltage which buffer 43 later transfers to the measuring 
circuit. After the predetermined measuring time interval .DELTA.T.sub.3, 
the controller 28 opens switches SW23 and SW4, allowing the buffer 43 to 
access the voltage held on the measuring capacitor C22. This voltage is 
advanced through the buffer amplifier 43 and switch SW22 (which the 
controller 28 closes at the time it opens switches SW23 and SW4), and is 
sampled on capacitor C7 at the input of the filter 39. The controller 28 
holds switch SW22 closed for a time duration which is sufficient for delta 
modulator 25 to convert the signal to digital form. After converting the 
sample to digital form, the controller 28 opens switch SW22 and closes 
switch SW26 to charge measuring capacitor C5 for the next measurement 
cycle. 
In one embodiment of the invention, the controller 28 measures each 
physiological parameter twenty times per second. The time duration 
.DELTA.T.sub.1 is 25 nanoseconds and the voltage measured on capacitor C2 
represents a patient motion parameter. The time duration .DELTA.T.sub.2 is 
250 nanoseconds and the voltage measured on capacitor C4 represents a 
respiration parameter. The time duration .DELTA.T.sub.3 is 500 nanoseconds 
and the voltage measured on capacitor C5 represents stroke volume. The 
operations of filter 23, capacitor C1 and delta modulator 25 are the same 
in FIG. 3 and FIG. 6. 
The filter 23 preferably filters the impedance signal in such a manner as 
to favor the sensing of a cardiac hemodynamic signal component over other 
physiological and non-physiological signal components. To this end, filter 
23 filters the impedance signal using a two-pole filter with a center 
frequency of 1.4 Hz. The gain of filter 23 is reduced by a factor of two 
(6 dB) at frequencies of 0.2 Hz and 10 Hz. 
The filter 34 preferably filters the impedance signal in such a manner as 
to favor the sensing of a respiration signal component over other 
physiological and nonphysiological signal components. To this end, filter 
34 filters the impedance signal using a two-pole filter with a center 
frequency of 0.2 Hz. The gain of filter 34 is reduced by a factor of two 
(6 dB) at frequencies of 0.05 Hz and 0.8 Hz. The filter 39 preferably 
filters the impedance signal in such a manner as to enhance the sensing of 
a patient motion signal component. To this end, filter 39 filters the 
impedance signal using a two-pole filter with a center frequency of about 
6 Hz. The gain of filter 34 is reduced by a factor of two (6 dB) at 
frequencies of 5 Hz and 10 Hz. 
Referring to FIG. 7, another embodiment of the impedance measurement 
circuit 14, which employs an inductor or coil 6 as an impedance sensor, is 
shown. The impedance measurement circuits of FIG. 3 and FIG. 7 are 
identical, except that the circuit of FIG. 7 includes the coil 6 for 
sensing impedance, while the circuit of FIG. 3 senses impedance by 
generating current pulses on and measuring the resulting voltage from lead 
11. Coil 6 at its proximal end is electrically connected to the lead 11 
between the coupling capacitor C3 and the electrode 10. The distal end of 
the coil 6 may be connected to pacemaker ground. Alternatively, the distal 
end of the coil 6 may connect to other points within or outside of the 
circuit. For example, the coil may connect with body fluids of the 
patient. 
The magnitude of the inductance of coil 6, which may range from 10 nH to 1 
mH, is selected to provide a high degree of electrical coupling to the 
tissue. For example, if the inductance is too small, the electrical field 
energy generated by the coil will be too small to detect changes in 
impedance which relate to physiological phenomena. Furthermore, the 
magnitude of the inductance may be varied according to the range of 
frequencies transmitted into the tissue. For example, an inductance of 10 
nH may be employed when a measuring frequency of about 330 MHz is applied 
to the tissue and an inductance of 1 mH may be used for a measuring 
frequency of approximately 500 kHz. 
The characteristics of the interrogated field may be relevant in 
determining the physical size of inductor 6. A larger sized inductor may 
be used to interrogate a specific area, such as a heart valve, a 
particular blood vessel or a heart chamber. A smaller sized inductor may 
be used to measure impedance in a more general area or for measuring 
multiple parameters. Furthermore, the inductor may be selected such that 
its physical size matches the dimensions of the organ or structure to be 
interrogated. 
The coil 6 can be resonated by appropriate selection of the interrogating 
frequency and the inductive and capacitive reactance of the circuit to 
increase the circulating current, thereby enlarging the measured field and 
raising the sensitivity of the measurement. 
The proximal connection of coil 6 may be located at a header (not shown in 
FIG. 7) of the pacemaker 5. The header is that portion of the pacemaker 
where the conductor of lead 11 is inserted into a terminal within the body 
of the pacemaker 5. The coil 6 may be located either externally of the 
pacemaker case 30 at the header, or internally of the pacemaker case. 
FIG. 8 is an illustration which depicts an embodiment of the present 
invention which employs a coil 6 at the header 6a of the pacemaker 5. The 
coil 6 is oriented so that it radiates its electrical field toward the 
patient's pleural area. 
Alternatively, the proximal connection of the coil 6 may be located at the 
distal end of the lead 11, preferably near the tip electrode 10. FIG. 9 is 
an illustration which depicts an embodiment of the present invention which 
employs a coil 6 at the distal end of the lead 11. The coil 6 is directed 
so that it radiates its electrical field into the patient's heart. 
FIG. 10 is a graph which illustrates the voltage amplitude V(C2), in volts, 
of a respiration signal as a function of the source capacitance (the 
measuring capacitance C2), in nanofarads, of the impedance measuring 
circuit of FIG. 3. In particular, FIG. 10 shows, for a given capacitor C2 
discharge time (a pulse width of 250 ns), the relationship between the 
change in voltage on the measuring capacitor C2 with respect to the load 
impedance and the change in load impedance due to respiration. The purpose 
of FIG. 10 is to show the importance of matching components of the source 
impedance of an impedance measuring circuit to the load impedance of the 
body. The capacitive discharge circuit depicted in FIG. 3 operates best 
with no impedance between the measuring capacitor C2 and the lead 11. 
Unfortunately, in an implantable pacemaker, a coupling capacitor C3 is 
generally perceived to be a requirement to assure safety of the patient. 
Therefore, the value of the measuring capacitor C2 is selected to best 
match the source impedance, which includes the measuring capacitor C2, 
with the load impedance, which includes the impedance of the coupling 
capacitor C3 and the impedance of the lead 11, in combination with the 
impedance of the body. The impedance measurement from the FIG. 3 
capacitive discharge circuit is derived from a direct measurement of 
voltage across a measuring capacitor C2 as a function of load resistance, 
the measuring capacitance and the initial voltage across the capacitor. 
For this capacitive discharge circuit, there exists a preferred measuring 
capacitance C2 for a given measuring pulse width (for example, 250 ns) and 
load impedance which will produce a maximum signal voltage V(C2). If the 
capacitance of the measuring capacitor C2 is very small in comparison to 
the load impedance of the body, the amplitude of the respiration signal is 
very small, leading to a modest signal to noise ratio and difficulty in 
appropriately controlling pacing rate. In contrast, FIG. 10 shows that 
values of capacitance of the measuring capacitor C2 which are large with 
respect to load impedance do not greatly diminish the respiratory 
impedance signal. Therefore, the capacitance of measuring capacitor C2 of 
an impedance measuring circuit should be equal to or larger than the 
capacitance which produces a maximum expected respiration signal. 
The graph of FIG. 11 characterizes the relative levels of different 
physiological and non-physiological signals which are detected by the 
circuit of FIG. 3 when it interrogates a patient's body with current 
pulses of different widths. It illustrates an important advantage of the 
pacemaker 5 of the present invention. The pacemaker can "tune" the 
impedance sensor to measure a particular type of signal and reject 
unwanted signals and other noise by selecting a particular measuring 
current pulse width. At very short pulse widths (e.g., 60 to 200 
nanoseconds) motion artifact signals have the largest amplitude, as shown 
by "motion" curve 35. The amplitude of physiological signals arising from 
the heart steadily rises with increasing pulse width duration, as shown by 
"cardiac" curve 36. The amplitude of respiratory signals abruptly rises 
with increasing pulse duration to pulse widths of about 250 ns, then 
decreases for larger pulse width durations, as shown by "respiratory" 
curve 37. The minute ventilation-controlled metabolic demand pacemaker of 
the present invention seeks a preferred pulse width of about 250 ns, which 
provides the best respiratory signal to noise ratio, as is illustrated at 
38 in FIG. 12, which figure comprises a graph that illustrates the level 
of a desired respiratory signal of FIG. 11 relative to a combination of 
non-respiration "noise" signals of that figure. A pulse width of this 
duration (250 ns) lessens the influence of cardiac signal "noise", avoids 
interface electrolytic phenomena, but still reduces the influence of 
motion artifacts. 
FIGS. 11 and 12 exemplify how different pulse widths provide for 
differentiation of signals arising from various physiological and 
non-physiological origins. Similarly, FIG. 13 illustrates this phenomenon 
in a sensing system which employs sinusoidal-like oscillating current 
modulation rather than discrete current pulses. Shorter pulse widths in a 
pulsed system have a similar effect upon signal sensing as higher 
frequencies in a sinusoidal-like oscillating current system. In general, 
the pacemaker 5 provides the best respiration signal sensing, in 
comparison with cardiac and motion noise, when the measuring current 
frequency is about 2 MHz. At higher frequencies, motion artifacts are 
large and at lower frequencies, cardiac signals obscure the respiration 
signal. 
The graph of FIG. 13 illustrates the signal amplitude arising from various 
physiological and non-physiological sources as a function of measuring 
sinusoidal-like oscillating current frequency. "Motion" curve 35a 
represents the amplitude of motion artifact signals; "cardiac" curve 36a 
represents the amplitude of physiological signals arising from the heart; 
and "respiration" curve 37a represents the amplitude of respiratory 
signals. The pacemaker may deliver these oscillating measuring currents in 
the form of sinusoidal-like oscillating waves or in the form of timed 
pulses of oscillating waves. Timed pulses of oscillating waves are 
discontinuous bursts of oscillating waves which last for a predetermined 
duration. For example, respiration may be measured by applying a 10 MHz 
oscillating wave burst lasting a duration of 100 ms. The measurements 
resulting from both methods are practically the same. To provide timed 
pulses of sinusoidal-like oscillating measuring current, the pacemaker 
deactivates the oscillating current to conserve energy, allow sensing of 
intracardiac electrograms or provide for generation of pacing pulses. The 
duration of timed pulses of sinusoidal-like oscillating current may range 
from one cycle of the oscillating frequency to essentially an infinite 
duration. 
The impedance measurement block 14 of FIG. 1 derives digital impedance 
samples, in the form of 8-bit data bytes having values ranging from -128 
to +127, at a rate of 20 per second and communicates these samples to the 
controller 28 by means of conductor 22. Negative digital signals carried 
by the conductor 22 indicate that the analog respiration signal is 
decreasing, while positive digital signals signify an increasing signal. 
Referring now to FIG. 14, wherein the circuit blocks of controller 28 are 
shown in greater detail, the manner in which minute volume is derived from 
the digital samples provided by impedance block 14 will now be considered. 
An absolute magnitude extractor 40 derives the absolute magnitude of each 
digital sample (i.e., negatively signed samples are changed to positive 
samples of the same amplitude). The average value of the digital samples 
is zero because the filter 23 (FIG. 3) in the impedance measurement block 
14 has a gain of zero for a DC input. By eliminating the sign from all 
samples, an averager 42 derives a running average of the absolute 
magnitudes of the samples. The time constant of the averager is short 
(e.g. about 25 seconds) so that the digital value at its output represents 
the average respiratory tidal volume over a few breaths. The absolute 
magnitude value of each sample represents the respiratory impedance 
signal. Therefore, the controller 28 adds and averages a sequence of these 
absolute magnitude sample values to provide a measure of the respiratory 
tidal volume. 
A sign extractor 44 monitors only the signs, and not the magnitudes, of the 
digital samples on conductor 22 to provide for zero crossing detection. 
The sign extractor 44 delivers successive bits, each of which represents 
the sign of a digital sample, to a zero crossing detector 46. The zero 
crossing detector 46 monitors respiration rate by ascertaining the timing 
of changes in the polarity of impedance measurement signal. Generally, a 
zero crossing occurs whenever the sign of a digital sample differs from 
the sign of the immediately preceding digital sample. However, there are 
physiological limits to respiration rate and, therefore, to the frequency 
of zero crossings. Zero crossings occurring at a rate higher than a 
predetermined physiological limit must indicate the presence of a noisy 
respiration signal. Thus, the zero crossing detector analyzes the signs of 
a number (for example, 10) of the most recently acquired samples and 
determines whether a defined preponderance of samples (for example, 7 of 
10) have a particular sign. If so, and if the last zero crossing operation 
which found a preponderance of a particular sign determined that the 
majority had an opposite sign, the zero crossing detector 46 presumes the 
occurrence of a zero crossing. When the sign changes, the zero crossing 
detector 46 triggers a sampler 48 to read the average value represented by 
the current value presented by the averager 42. The sampler 42 delivers 
this average value to both a short-term averager 50 and a long-term 
averager 52. In the preferred embodiment of the invention, the short-term 
averager 50 has a time constant of slightly less than a minute and the 
long-term averager 52 has a time constant of about one hour. 
The zero crossing detector 46 pulses its output twice, and the sampler 48 
samples twice, during each breath, when the impedance signal crosses zero 
during exhalation and during inhalation. The zero crossing detector 46 
employs the previously described "majority vote" technique to sense a zero 
crossing, in which the detector assumes an occurrence of a zero crossing 
when a predetermined proportion of the most recent samples have a sign 
opposite to that of the sign determined after the last zero crossing. In 
the preferred embodiment of the invention, at least 70% of the most recent 
samples in the last 0.5 second must have a sign opposite to that of the 
sign determined after the last zero crossing. 
Each average value sample at the output of averager 42 represents the tidal 
volume, the average of the last few integrals of the respiratory impedance 
signal. The short-term averager 50 and the long-term averager 52 derive 
values which are dependent not only on the magnitudes of the samples, but 
also upon the rate of the oscillating respiratory signal, as determined by 
the zero crossing detector 46. Because the long-term and short-term 
averagers update and accumulate samples at each zero crossing event, the 
long-term and short-term minute volume values reflect the rate of 
breathing as well as the depth of breathing. 
As shown in FIG. 14, a summer 54 derives .DELTA.MV, the difference between 
the short-term averaged and long-term averaged minute volume signals. 
.DELTA.MV is the control signal which drives the pacing rate. As the 
short-term average increases relative to the long-term average, 
representing an increasing metabolic demand, the pacing rate increases. 
Conversely, when .DELTA.MV decreases, the pacing rate decreases. 
The .DELTA.MV value at any instant is the input to a limiter 56, which 
compares .DELTA.MV to .DELTA.MVMAX, a predetermined value which serves as 
the maximum .DELTA.MV value allowed to control the pacing rate. The 
limiter 56 applies the current value of .DELTA.MV, or .DELTA.MVMAX if it 
is smaller than .DELTA.MV, to the minus input of a summer 58. The summer 
58 compares the output of limiter 56 to "maximum interval", a quantity 
applied at the plus input of summer 58 which represents an offset 
corresponding to a physician-determined minimum pacing rate. Summer 58 
continuously presents its output, a difference value called "respiration" 
pacing rate 96, to the input of a rate control block 95, which compares 
and analyzes derived pacing rates, for example "cardiac" rate 98 and 
"activity" rate 97 as determined by other physiological signal processors 
which are to be described hereinafter, to determine a pacing rate for 
stimulating the patient's heart. The rate control block 95 uses this 
pacing rate, in combination with a SENSE signal 24 from the sense 
amplifier 16 of FIG. 1 to determine when to deliver a pace signal 26 to 
the pulse generator 18 of FIG. 1. 
In this manner, the pacemaker operates in a standard VVI mode except that 
the minute volume measurement may determine the pacing rate. As the 
quantity .DELTA.MV increases, the summer 58 derives a smaller difference 
value ("maximum interval" minus .DELTA.MV) that it presents to rate 
control block 95. This, in turn, means that the pacing rate increases, as 
is required for a larger .DELTA.MV. When .DELTA.MV is zero, the summer 58 
presents the "maximum interval" value to the rate control block 95, which 
results in the minimum pacing rate, precisely what is required when there 
is no metabolic demand beyond that provided by the minimum pacing rate. 
The quantity "maximum interval" is simply the interval which corresponds 
to the minimum rate. 
Timing for the rate control block 95 is provided by a clock 64 which 
applies pulses to a divider 66. The divider 66 divides the clock pulses by 
a quantity referred to as a prescaler and produces a count signal on line 
62 which extends to the rate control block 95. 
Conventional pacemakers include telemetry systems, as represented by block 
68 of FIG. 14, which allow a physician to program parameters such as 
minimum rate, as well as the prescaler value, .DELTA.MVMAX and the 
reference threshold which is applied to a comparator 60, as will be 
described below. The method of derivation of these programmable parameters 
is disclosed in the description of the '725 pacemaker. 
Again referring to FIG. 14, the output of summer 54 is input, not only to 
the limiter 56, but also to the plus input of comparator 60. The 
telemetrically-programmed reference threshold feeds the minus input of the 
comparator. Whenever .DELTA.MV exceeds the reference threshold, the output 
of the comparator goes high and inhibits the long-term averager 52. In 
effect, a large value of .DELTA.MV represents a metabolic demand which is 
associated with an exercising patient. Until the patient stops exercising, 
the long-term average does not increase. If it were allowed to increase, 
after an hour or more the long-term average would approach the value of 
the short-term average, .DELTA.MV would diminish and the pacing rate would 
drop from its original high value. Once the patient begins exercising and 
the pacing rate increases, it is not desirable that the rate decrease 
simply due to the elapse of time. For this reason, the pacemaker fixes the 
long-term average. When the patient stops exercising and the short-term 
average decreases, .DELTA.MV will fall below the reference threshold and 
the long-term average will again track the short-term average in the usual 
manner. In the illustrative embodiment of the invention, the reference 
threshold is equal to one-half of the .DELTA.MVMAX, unless the physician 
programs the value differently. This technique allows long-term adaptation 
to a basal minute volume measurement level while still allowing extended 
periods of exercise. 
Referring now to FIG. 15, wherein the functional blocks, designated 
generally as an activity signal processor 79, performed by controller 28 
are shown in greater detail. The flow chart of FIG. 15 illustrates the 
manner in which the controller 28 derives the value of patient motion, or 
activity, from the digital samples 22 provided by impedance block 14 of 
FIGS. 3, 6 or 7. The activity signal processor 79 reads digital samples 22 
and applies them to a bandpass filter 80 which rejects low and high 
frequency components of the digital signal. Next, a threshold detect block 
82 compares the filtered signal amplitude to a preset threshold value. 
Under the control of threshold logic block 84 and a maintain "activity" 
rate block 86, if the signal does not surpass the threshold value, an 
"activity" pacing rate 97 is set to a predetermined baseline rate and sent 
to the rate controller 95, which comprises another group of function 
blocks that are performed by the controller 28. If the filtered signal is 
larger than the threshold value, averager block 88 sums or digitally 
integrates the digital sample value over a selected time period to 
determine a time average of the activity signal. Next the averaged signal 
extends to a slope mapper 90 which converts the current processed activity 
sample value to a pacing rate value, according to a predetermined and 
selected linear or nonlinear slope relationship. A rate limiter block 92 
compares this pacing rate value to preselected upper and lower limits and, 
if the rate is outside the limits, will set the pacing rate to the 
appropriate limit. 
Next, a determine "activity" rate block 94 sets the pacing rate which is 
determined according to the activity measurement. Block 94 may set the 
"activity" rate 97 to the rate from the rate limiter 92 or may provide 
rate smoothing or averaging of the "activity" rate if the difference 
between the current and most recent rates differ by more than a 
predetermined amount. 
Telemetric programming by a physician may be used to program the selected 
time period, slope, upper and lower pacing rates and rate smoothing 
variables. 
Referring now to FIG. 16, wherein the functional blocks, designated 
generally as a cardiac signal processor 112, performed by controller 28 
are shown in greater detail. The flow chart illustrates the manner in 
which the controller 28 analyzes the digital samples 22 provided by the 
impedance block 14 of FIGS. 3, 6 or 7 and derives the value of cardiac 
rate signal 98 which is indicative of the patient's stroke volume. 
Referring to FIG. 1 in conjunction with FIG. 16, in wait for cardiac 
signal block 100, the controller 28 waits for the occurrence of either a 
E signal on line 26 or a SENSE signal on line 24, and then, in either 
case, waits a predetermined delay time .tau..sub.1 before performing 
bandpass filter block 102. FIG. 17A is an illustration of an intracardiac 
electrogram signal portraying an intrinsic cardiac signal which evokes a 
SENSE signal on line 24. FIG. 17B is an illustration of a cardiac 
hemodynamic signal which results from the heart depolarization 
characterized in the intracardiac electrogram of FIG. 17A. During wait for 
cardiac signal block 100, the controller waits for the delay period 
.tau..sub.1 to expire, as is shown in FIG. 17B. The controller 28 samples 
the cardiac signal only for the time .tau..sub.2 when a hemodynamic signal 
is present following a cardiac electrical event (E or SENSE). The 
controller 28 sets the predetermined delay time .tau..sub.1 to correspond 
to the time between the cardiac electrical event and the mechanical 
ejection of blood from the heart. The delay following a SENSE signal may 
be different from the delay after a E signal. 
Following the triggering signal and delay period, the cardiac signal 
processor 112 (FIG. 16) reads digital samples 22 and applies them to a 
bandpass filter 102 which rejects low and high frequency components of the 
digital signal. Averager block 104 sums or digitally integrates the 
digital sample value over a selected time period .tau..sub.2, which is 
illustrated in FIG. 17B, to determine a time average of the stroke volume 
signal. During the time interval .tau..sub.2, the controller 28 controls 
the sampling rate of the digital samples. A preferred sampling rate is 
approximately 100 Hz, although the controller 28 regulates sampling rates 
in the range from 20 Hz to 200 Hz. The preferred sampling time interval 
.tau..sub.2 is 200 milliseconds, although the controller 28 confines this 
interval to range from 25 milliseconds to 250 milliseconds. 
Averager block 104 counts the number of samples processed in bandpass 
filter block 102 and averager block 104 to correlate the same with the 
stroke volume signal time period .tau..sub.2. If more samples are to be 
processed in this cardiac cycle, more samples logic block 105 returns 
control of the procedure to bandpass filter block 102 to process another 
sample. Otherwise, the cardiac signal processor 112 performs the stroke 
volume to cardiac rate converter 106 operation. The averager block 104 
provides any required time and amplitude scaling to appropriately 
correlate the cardiac and respiration signals. The stroke volume to 
cardiac rate converter 106 converts the current processed stroke volume 
sample value to a rate. 
In the preferred embodiment of the invention, the stroke volume to heart 
rate converter 106 determines heart rate from the stroke volume 
measurement by one of three methods. A physician may select a desired one 
of such methods by telemetric programming. According to a first method, 
the stroke volume to heart rate converter 106 derives the pacing rate by 
determining a value for the average stroke volume at rest, S.sub.Ravg, 
through time averaging of stroke volume values which have been accumulated 
only when the cardiac rate, either sensed or stimulated, is below a 
predetermined resting rate, HR.sub.rest. The stroke volume to heart rate 
converter 106 continuously derives a short-term average stroke volume, 
S.sub.ST, by averaging stroke volume values, regardless of heart rate, 
only over a short time (for example, 20 seconds). The stroke volume to 
heart rate converter 106 then subtracts the average stroke volume at rest 
(S.sub.Ravg) from the short-term average stroke volume, S.sub.ST, to 
determine the increment in stroke volume, S.sub.inc. The stroke volume to 
heart rate converter 106 sets S.sub.inc to zero if it is otherwise 
negative. In general, an increasing metabolic demand causes an elevation 
in stroke volume, but an increasing heart rate from pacing causes a 
decrease in stroke volume. Therefore, the stroke volume to heart rate 
converter 106 increases pacing rate as the increment in stroke volume, 
S.sub.inc, increases until an increment in pacing rate causes a reduction 
in S.sub.inc. When this occurs, the stoke volume to heart rate converter 
106 lessens the pacing rate by a predefined rate decrement and holds the 
pacing rate at this level for a predetermined length of time before 
further raising the pacing rate. 
According to a second method, the stroke volume to heart rate converter 106 
derives the pacing rate by determining a value for the average cardiac 
output at rest, CO.sub.Ravg. According to this method, the stroke volume 
to heart rate converter 106 multiplies stroke volume values by heart rate 
to determine an instantaneous cardiac output value, CO. CO samples are 
accumulated to provide CO.sub.Ravg only when the cardiac rate, either 
sensed or stimulated, is below the predetermined resting rate, 
HR.sub.rest. The stroke volume to heart rate converter 106 continuously 
derives a short-term average cardiac output, CO.sub.ST, by averaging CO 
values, regardless of heart rate, only over a short time (for example, 20 
seconds). The stroke volume to heart rate converter 106 then subtracts the 
average cardiac output at rest (CO.sub.Ravg) from the short-term average 
cardiac output, CO.sub.ST, to determine the increment in cardiac output, 
CO.sub.inc. The stroke volume to heart rate converter 106 sets CO.sub.inc 
to zero if it is otherwise negative. Cardiac output must increase with 
pacing rate, otherwise the pacing rate should no longer increase. 
Therefore, the stroke volume to heart rate converter 106 increases pacing 
rate as the increment in cardiac output, CO.sub.inc, increases until an 
increment in pacing rate causes a reduction in CO.sub.inc. When this 
occurs, the stoke volume to heart rate converter 106 reduces the pacing 
rate by a predefined rate decrement and holds the pacing rate at this 
level for a predetermined length of time before further raising the pacing 
rate. 
According to a third method, the stroke volume to heart rate converter 106 
does not independently change the pacing rate, but rather allows the 
pacing rate to be determined by the other rate processors, either the 
activity signal processor or the respiration signal processor. In this 
method, the stroke volume to heart rate converter 106 continuously 
determines the increment in cardiac output, CO.sub.inc, in the manner of 
the second method described above. If the pacing rate, as determined by 
the non-cardiac signal processor, increments the pacing rate while 
CO.sub.inc is decreasing, the stroke volume to heart rate converter 106 
will decrease the pacing rate by a predefined rate decrement and holds the 
pacing rate at this level for a predetermined length of time before 
further raising the pacing rate. 
Assuming that the stroke volume to heart rate converter 106 sets a pacing 
rate in accordance with one of the methods described above, a rate limiter 
block 108 compares this pacing rate value to preselected upper and lower 
limits and, if the rate is outside the limits, will set the "cardiac" 
pacing rate to the appropriate limit. 
Next, a determine "cardiac" rate block 110 sets the pacing rate which is 
determined according to the stroke volume measurement. Block 110 may set 
the "cardiac" rate to the rate from the rate limiter 108 or may provide 
rate smoothing or averaging of the "cardiac" rate if the difference 
between the current and most recent rates differ by more than a 
predetermined amount. 
As in the case of the respiration and activity signal processors, 
telemetric programming by a physician may be used to program the selected 
time period, slope, upper and lower pacing rates and rate smoothing 
variables. 
Referring to FIG. 18, wherein the functional blocks performed by rate 
control block 95 of FIG. 14 are shown in greater detail, rate control 
block 95 analyzes the pacing rates from the various physiological signal 
processors to determine an operational pacing rate. A threshold block 114 
reads the "activity" rate 97, "respiration" rate 96 and "cardiac" rate 98 
inputs and compares each rate with a corresponding activity, respiration 
and cardiac rate threshold. Under the control of logic block 116, if none 
of the rates are above the corresponding threshold value, a compare 
present rate to baseline rate block 118 compares the present rate to a 
predetermined and preprogrammed baseline pacing rate. Under the control of 
logic block 120, if the present rate is equal to the baseline rate, a set 
rate to baseline block 122 sets the pacing rate for the next cardiac cycle 
to the baseline rate. Block 122 (as well as blocks 124, 130, 132 and 138, 
which are described below) includes access to a timer (not shown) which 
the controller 28 loads with the most recently determined operational 
pacing rate. This operation takes place upon the occurrence of either of 
two events, a SENSE signal upon conductor 24 or the countdown to zero of 
timer 62. 
In this manner the count 62 pulses, which act upon the decrement input of 
the timer, occur at a rate slower than the clock rate. The count in the 
timer decrements whenever a pulse appears at the output of divider 66 in 
FIG. 14. When the timer decrements to zero, it produces a pulse upon E 
conductor 26 to trigger activity of the pulse generator 18 in FIG. 1. If 
the sense amplifier 16 of FIG. 1 senses a natural heartbeat before the 
timer decrements to zero, the timer will not produce a pulse on E 
conductor 26. In either case, the timer loads the pacing rate which is set 
within rate control block 95 to initialize the escape interval of the 
pacemaker. The escape interval is the time between a paced or sensed 
cardiac event and the subsequent pacing stimulus. If the present rate is 
not equal to the baseline rate, smooth rate to baseline block 124 may 
determine the difference between the present rate and the baseline rate 
and perform rate smoothing to gradually decrement the pacing rate to the 
baseline level. Programming of the rate control block 95 may also 
designate that no rate smoothing is to occur, in which case, block 124 
will set the pacing rate to the baseline level. Block 124 acts in the 
manner of block 122 to operate the timer and generate pace 26 signals 
which activate the pulse generator 18. 
Referring back to logic block 116, if one or more of the physiological 
pacing rates did exceed its corresponding threshold rate, logic block 125 
determines whether the activity rate was the only parameter above its 
threshold rate. The rate control block 95 performs an extra check on the 
activity parameter because it is inherently susceptible to activation by 
noisy or nonphysiological signals. If activity is the only parameter 
providing a rate which is above threshold, how long activity rate setting 
block 126 determines how long activity has been the only operative signal. 
Under the control of logic block 128, if this condition has persisted for 
longer than a predetermined duration, smooth rate to baseline block 132 
smoothes (averages) the pacing rate to the baseline rate in the manner of 
block 124. If the activity rate has not been the only operative signal for 
the prescribed duration, set rate to activity rate block 130 employs the 
"activity" rate to set the operational pacing rate and generate pace 26 
signals for the page generator 18 in the manner of block 122. It is 
desirable to allow the activity physiological parameter to drive the 
pacing rate because its sensor reacts to stimulation very rapidly. 
However, that the activity signal persists while no other sensor responds 
is an indication of no true physiological stimulation. 
Also under the control of logic block 125, if any sensor other than the 
activity sensor produces a rate above its threshold rate, block 134 
compares the derived "activity", "respiration" and "cardiac" rates. Block 
136 sets the pacing rate to the fastest of these rates. Block 138 then 
issues a pace 26 signal to the pace generator 18 and may be programmed to 
gradually smooth the pacing rate from its current level to its derived 
level. 
From the foregoing discussion, it is apparent that the present invention 
provides rate-responsive pacing based on measurements of respiratory 
minute volume, patient motion or activity, and cardiac mechanical signals 
corresponding to stroke volume which are sensed from a standard pacing 
lead using a single sensor. The invention accomplishes substantial 
improvements in the determination of a physiological pacing rate by 
employing the sensing of multiple physiological parameters using a single 
impedance transducer. 
Although the invention has been described with reference to particular 
embodiments, it is to be understood that such embodiments are merely 
illustrative of the application of the principles of the invention. 
Numerous modifications may be made therein and other arrangements may be 
devised without departing from the true spirit and scope of the invention.