Physiological occurrence, such as apnea, monitor and X-ray triggering device

A respiration monitor and X-ray trigger apparatus wherein signals are generated indicative of respiration rate, respiration extrema, and heart rate. A trigger signal is selectively generated to the X-ray machine just prior to occurrence of a selected respiration extrema. The value of the respiration signal at a sample occuring just after detection of the selected extrema is stored, and modified in accordance with the resolution of the system. The X-ray trigger signal is generated, if not otherwise inhibited, upon detection of the next occurrence of the modified value in the respiration wave form. Apnea alarms are generated if a predetermined decrease in respiration rate is detected in successive periods in conjunction with a predetermined number of decelerating heart beats. An alarm is also generated if the respiration rate and heart rate signals appear to indicate that the respiration and heart rates are equal for a predetermined number of periods, or if the respiration or heart rates stray outside of preset threshold values. Provisions are included for adaptive DC offset, and to eliminate electromagnetic interference. Amplitude, time domain, and slew rate discrimination against artifacts are also provided.

The present invention is directed to a respiration monitor, and, in 
particular, to an apparatus for triggering an X-ray machine at a 
predetermined desired point in the respiration cycle. 
Respiration monitors, are, in general, well known. The greater percentage 
of the respiration monitors are, more accurately described as apnea 
monitors. Apnea is classically defined as the cessation of breathing 
efforts. Other monitors provide indicia of the respiration rate of the 
subject. 
Various methods of sensing respiration have been utilized. For example, the 
air exchange between the subject's lungs and an outside source has been 
directly measured. A thermistor anemometer provides an electrical signal 
indicative of air flow by placing a thermistor in a face mask, by the 
nares, or mouth or in the pharynx. The resistance of the thermistor 
changes as it is warmed and cooled by the subject's breathing. For further 
description of such systems, reference is made to Gordon, D. H., Thompson, 
W. L.: "A new Technique for Monitoring Spontaneous Respiration", Med 
Instrum, Vol. 9, Jan/Feb. 1975, pp. 21-22; Gulleminault, C. et al: 
"Abnormal Polygraphic Findings in Near-Miss Sudden Infant Death", The 
Lancet, Vol. 1, June 19, 1976, p. 1326; HEALTH DEVICES, Evaluation: 
Infant Apnea Monitors, Vol. 4, Nov. 1974, p. 3; Lipton, E. L. et al: "A 
Respiratory Alarm for Infants", J. Pediatr., Vol. 65, Aug. 1964, pp. 
294-296; Pope, J. M., Dimeff, J., Abraham, S.: "A Wireless Respiration 
Failure Detection System", Med. Biol. Eng., May 1974, pp. 348-354; and 
Rigatto, H., Brady, J. P.: "A New Nosepiece for Measuring Ventilation in 
Pre-term Infants", J. Appl. Physiol., Vol. 32, March 1972, pp. 423-424. 
Other methods, involving the subject breathing into a face mask include 
the reverse plethysmograph (wherein volume changes in a reservoir is 
measured with a strain gauge), a spiro meter (wherein the volume of air 
exhaled into a reservoir is measured) and a wet-test meter. For more 
complete descriptions of these respiration sensing techniques reference is 
made to Avery, M. E., O'Doherty, N.: "Effects of Body Tilting on the 
Resting End-Expiratory Position of Newborn Infants", Pediatrics, Vol. 29, 
Feb. 1962, pp. 255-260 and Miller, H. C., Snull, N. W.: "Further Studies 
on the Effects of Hypoxia on the Respiration of Newborn Infants", 
Pediatrics, Vol. 16, 1955, pp. 93-103. A pneomotachograph, that is a fine 
mesh screen which senses pressure drops has also been used to sense 
respiration. Reference in this regard is made to Polgar, G.: "Comparison 
of Methods for Recording Respiration in Newborn Infants", Pediatrics, Vol. 
36, Dec. 1965, pp. 861-868; Stark, A. R., Thach. B. T.: "Mechanisms of 
Airway Obstruction Leading to Apnea in Newborn Infants", J. Pediatr., Vol. 
89, Dec. 1976, pp. 982-985; and Warburton, D., Stark, A. R., Taeusch, H. 
W.: "Apnea Monitor Failure In Infants With Upper Airway Obstruction", 
Pediatrics, Vol. 60, Nov. 1977, pp. 742-744. 
Other direct respiration sensing techniques include monitoring the 
biometric pressure changes in a closed chamber within which the subject is 
disposed, (see Polgar, G.: "Comparison of Methods for Recording 
Respiration in Newborn Infants", Pediatrics, Vol. 36, December 1965, pp. 
861-868); measurement of the partial pressure of carbon dioxide (CO.sub.2) 
in the subject's exhilation (end-tidal PCO.sub.2) and disposing a 
microphone on the subject's neck to sense the sound of air movement in the 
throat has been utilized. 
Other methods provide signals indicative of respiration by sensing body 
motion. For example, air mattresses (wherein movement displaces air to 
cool a heated thermistor), air filled with rubber vests, capacitance 
pneumography (wherein the subject lies between two conductive plates such 
that breathing movements change the capacitance between the plates), 
capacitance mattresses, displacement magnitometers, displacement 
transducers, electret pressure transducers, permanent magnetic motion 
transducers, piezo resistive mattresses, radar movement sensors, 
ultrasonic movement sensors, and strain gauges have been utilized to 
provide signals indicative of respiration. Further description of 
capacitance pneumography is provided in Barrow, R. E., Colgan, F. J.: "A 
Noninvasive Method for Measuring Newborn Respiration", Respir. Care, Vol. 
18, 1973, pp. 412-414; Franks, C. I., Brown, B. H., Johnston, D. M.: 
"Contactless Respiration Monitoring of Infants", Med. Biol. Eng., Vol. 14, 
May 1976, pp. 306-312 and Sigdell, J. E.: "A Theoretical Study of 
Capacitive Plethysomography", Med. Biol. Eng., Vol. 9, Sept. 1971, pp. 
447-457. Description of displacement magnetometers are provided in Knill, 
R. et al: "Respiration Load Compensation in Infants", J. Appl. Physiol., 
Vol. 40, March 1976, pp. 357-361 and Mead, J. et al: "Pulmonary 
Ventilation Measured From Body Surface Movements", Science, Vol. 156, 
1957, pp. 1383-1384. The electret pressure transducer and permanent 
magnets motion sensors are described in Franks, C. I., Brown, B. H., 
Johnston, D. M.: "Contactless Respiration Monitoring of Infants", Med. 
Biol. Eng., Vol. 14, May 1976, pp. 306-312 and HEALTH DEVICES, Evaluation: 
Infant Apnea Monitors, Vol. 4, Nov. 1974, pg. 3. For a description of a 
suitable piezo resistive mattress, reference is made to Smith, J. E., 
Scopes, J. W.: "A New Aponea Alarm for Babies", Lancet, Vol. 2, 1972, pp. 
545-546. Strain gauge techniques are described in Paulev, P. E. et al: 
"Strain-Gauge Versus Water Plethysmography: Description of Simplified 
Systems and Analysis of Differences and Accuracy", Med. Biol. Eng., Vol. 
12, July 1974, pp. 437-445 and Polgar, G.: "Comparison of Methods for 
Recording Respiration in Newborn Infants", Pediatrics, Vol. 36, December 
1965, pp. 861-868. Radar sensing techniques are described in Bloice, J. 
A., Caro, C. G.: "Contactless Apnoea Detector Using Low Energy Radar" , J. 
Physiol., Vol. 223, Feb. 1972, pp. 3-4; Caro, C. G., Bloice, J. A.: 
"Contactless Apnoea Detector Based on Radar", Lancet, Vol. 2, Oct, 30, 
1971, pp. 959-961 and Franks, C. I., Brown, B. H., Johnston, D. M.: 
"Contactless Respiration Monitoring of Infants", Med. Biol. Eng., Vol. 14, 
May 1976, pp. 306-312. And in U. S. Pat. No. 3,993,995 (Kaplan et al., 
1976). Ultrasonic motion sensors are described in "Ultrasonic Recording of 
Fetal Breathing", H. B. Meire et al, British Journal of Radiology, Vol. 
48, No. 570, pp. 477-480, January 1975, patent associated literature 
B010-7506-B and U.S. Pat. No. 3,802,253 (Lee, 1974), U.S. Pat. No. 
3,856,985 (Yokoi et al, 1974) and U.S. Pat. No. 3,864,660 (Ranalli et al, 
1975). 
Other respiration sensing techniques include the use of an electromyograph, 
which senses electrical activity associated nerve and muscle activation 
during breathing; use of a catheter placed in the esophagus to measure 
intra-esophageal pressure (reflecting intrathoracic pressure) during 
breathing; and use of photoelectric plethysomograph (hemodensitomer) which 
senses the absorption of particular wavelengths of incident light by the 
subject's blood. For a description of intra-esophageal pressure 
measurement techniques reference is made to Stark, A. R., Thach, B. T.: 
"Mechanisms of Airway Obstruction Leading to Apnea in Newborn Infants", J. 
Pediatr., Vol. 89, December 1976, pp. 982-985 and with respect to the 
hemodensitomer, Wallace, J. D., et al: "Observation of Pulse and 
Respiration in the Neonate, A Preliminary Report", IEEE Trans. Biomed. 
Eng., Vol. 20, Sept. 1973, pp. 388-391. 
Still another respiration sensing technique is that of impedance 
pneumography. Impedance pneumography involves sensing the changes in 
impedance across the thorax during respiration. A 10 kHz-100 kHz electric 
current on the order of 50 .mu.a is passed through electrodes taped on the 
subject's thorax. A voltage is developed and measured. The voltage, 
proportional to the transthoracic impedance, varies directly in accordance 
with the respiratory cycle. The change in impedance is believed to be 
caused by a change in blood volume in the vessels supplying blood to the 
lungs and in part to the volume of air in the lungs themselves. 
All of the above noted sensing techniques provide electrical signals which 
are directly indicative of the respiration cycle. The respiration cycle is 
then analyzed and apneic episodes detected, respiratory rate computed, or 
both. 
Other apnea monitors sense various short term physiological effects of 
apnea. When respiration ceases, changes in the partial pressure of 
arterial oxygen (P.sub.AO.sbsb.2), arterial CO.sub.2 (P.sub.AOC.sbsb.2), 
arterial pH, heart rate, skin color and blood pressure are evidenced. For 
example, heart rate generally decreases during apnea. Some intensive care 
units utilized heart rate monitors rather than respiration monitors to 
detect apnea. 
Use of apnea monitors is particularly wide spread in infant intensive care 
units. Newborn infants with respiratory problems are particularly subject 
to apneic episodes. Often only a mild stimulus such as a noise or light 
body contact (e.g., squeezing the infants foot) is necessary to induce the 
infant to resume breathing. In other instances, only a vigorous shake or 
body repositioning is necessary to induce resumption of breathing. In 
other cases, use of resuscitator mechanisms are necessary to reinitiate 
breathing. However, while resumption of breathing may often easily be 
induced, irreversible damage to the infant can occur if the apneic episode 
is not promptly detected. Apnea undetected for as little as 30 seconds can 
cause irreversible damage to the infant and can be fatal if undetected for 
as little as a minute. 
The problems with presently available commercial respiration monitors and 
apnea detectors are addressed in "Evaluation Infant Apnea Monitors", 
HEALTH DEVICES, Vol. 4, No. 1, 1974 by the Emergency Care Research 
Institute, and in the final report "An Investigation to Determine the 
Risks and Hazards Associated with Apnea Monitors", E.C.R.I. 410-276, July 
1978 presented to the Bureau of Medical Devices of the Federal Food and 
Drug Administration. In the final report, it is noted that the primary 
shortcoming of the presently available apnea monitors is that the monitors 
tend to fail to discriminate between artifacts and the actual respiration 
waveform. 
Commercially available apnea detectors are particularly subject to 
detection of signals not related to breathing (artifacts) and erroneously 
identifying the artifact signals as the respiratory waveform. Artifacts 
commonly arise from cardiovascular activity, electromagnetic interference, 
self-noise, vibration, line voltage fluctuation, poor electrode contact 
and placement and neuromuscular activity. Detection of artifact signals as 
the respiration waveform is referred to as a "false negative". 
With respect to the problems of false negatives, reference is made to 
Blake, A. M. et al., "Clinical Assessment of Aponea-Alarm Mattress for 
Newborn Infants", Lancet, July 1970, pp. 183-185; Edwards, N. K. et al, 
"Phantom Breathing in Monitored Infants", Am. J. Dis. Child., Vol. 125, 
May 1973, pp. 684-685; Guilleminault, C. et al, "A Polygraphic Study of 
the Sleep and Respiration Patterns of Apneic Premature Infants On and Off 
an Oscillating Water Bed", In Apnea of Prematurity, Report of the 
Seventy-First Ross Conference on Pediatric Research, Lucey, J. F., 
Shannon, D. C., Soyka, L. F., eds, Columbus, Ohio, Ross Laboratories 1977, 
pp. 29-34; Guilleminault, C. et al, "Apneas During Sleep In Infants: 
Possible Relationship with Sudden Infant Death Syndrome", Science, Vol. 
190, Nov. 1975, p. 677; HEALTH DEVICES, Evaluation: Infant Apnea Monitors, 
Vol. 4, Nov. 1974, p. 3; Lewak, N., "Sudden Infant Death Syndrome in a 
Hospitalized Infant on an Apnea Monitor", Pediatrics, Vol. 56, 1975, pp. 
296-298; Mangat, D., Orr, W. C., Smith, R. P., "Sleep Apnea, 
Hypersomnolence, and Upper Airway Obstruction Secondary to Adenotonsillar 
Enlargement", Arch Otolarygol, Vol. 103, July 1977, pp. 383-386; Peabody, 
J. L., Phillip, A. G. S., Lucey, J. F., "Disorganized Breathing--an 
Important Form of Apnea and Cause of Hypoxia", Society for Pediatric 
Research Annual Meeting, San Francisco, CA, Apr. 27-29, 1977. Abstracted, 
Pediatr. Res. 11:540, 1977; Peabody, J. L. et al, "Failure of Conventional 
Respiratory Monitoring to Detect Hypoxia", Society for Pediatric Research, 
San Francisco, CA, Apr. 27-29, 1977. Abstracted, Pediatr. Res. 11:539, 
1977; Rolfe, P., "Monitoring in New Born Intensive Care". Biomed. Eng., 
Vol. 10, Nov. 1975, pp. 399-404; Stark, A. R. et al, "The Pediatric 
Pneumogram: A New Method for Detecting and Quantitating Apnea in Infants", 
J. Pediatr., Vol. 89, Dec. 1976, pp. 982-985; and Warburton, D. et al, 
"Apnea Monitor Failure in Infants With Upper Airway Obstruction", 
Pediatrics, Vol. 60, Nov. 1977, pp. 742-744. 
It has been further noted that available apnea monitors which utilize 
sensors which do not measure the actual exchange of air are inadequate in 
detecting upper airway obstructions. 
Conversely, a second shortcoming of the presently available apnea monitors 
is a tendency to trigger false apnea alarms while the infant is breathing 
normally (false positive). False positive alarms do not pose an immediate 
threat to the subject. However, such failures tend to undermine the 
operator's confidence in the machine, sometimes causing personnel to 
ignore or disable the alarms. In this regard, reference is made to 
Edwards, N. K. et al, "Phantom Breathing in Monitored Infants", Am. J. 
Dis. Child., Vol. 125, May 1973, pp. 684-685; Henning, L, "Impact of the 
Apnea Monitor on Family Life", NICHD, Prenatal Biology and Infant 
Mortality Branch, 1974, pp. 19-22; Rolfe, P., "Monitoring in New Born 
Intensive Care", Biomed. Eng., Vol. 10, Nov. 1975, pp. 399-404; and Stein, 
I. M. et al, "The Pediatric Pneumogram A New Method for Detecting and 
Quantitating Apnea in Infants", Pediatrics, Vol. 55, May 1975, pp. 
599-603. Systems using sensors which directly measure air exchange, 
however, are often not suitable for use on infants. For example, 
monitoring techniques using face masks obscure the view of the nose and 
mouth. The face mask in and of itself, tends to cause a resistance to 
breathing and an increase in anatomical dead space. The face masks are 
generally too large and cumbersome for use on infants and entail a 
likelihood of rebreathing. In addition, the masks are often difficult to 
clean. 
The various mattress and transducer pad monitoring techniques are extremely 
sensitive to body movements and vibrartion artifacts. The mattress sensors 
are often extremely susceptible to physical damage and deterioration and 
are hard to clean. 
Irrespective of the method of respiration sensing utilized, the respiration 
monitor must be compatible for use with other diagnostic machinery such as 
and ECG, X-ray machines or defibrillating apparatus. For example, where an 
infant is in an intensive care situation, respiration and ECG are normally 
monitored. If the infant is removed from the monitor in order to take 
X-rays, there is a clear and present danger that the infant will have an 
apneic episode or cardiac arrest during the period when the X-rays are 
being taken. In graphic terms, there is a likelihood that good X-rays 
pictures will be taken of a dead infant. Accordingly, it is desirable that 
all diagnostic analysis be provided by a single monitor, compatible with 
an X-ray machine. 
In addition, it is often desirable to utilize a respiration monitor for 
triggering an X-ray machine. For example, X-ray pictures for study of the 
lungs should be taken at points of maximum inhalation when the lungs are 
fully inflated. Similarly, X-ray pictures for examination of the area 
surrounding the lungs should be taken at points of maximum exhalation, 
when the lungs are fully deflated. Such instances of maximum inhalation or 
exhalation shall hereinafter be collectively termed "respiration extrema". 
X-ray pictures for examination of other thoracic areas are also taken at 
points of respiration extrema to minimize motion artifacts. However, 
triggering an X-ray machine at the proper time within the respiratory 
cycle is exceedingly difficult with respect to an uncooperative subject 
such as an infant. Examples of respiration monitor apparatus used for 
triggering an X-ray machine in synchronism with the respiratory cycle are 
described in the aforementioned U.S. Pat. Nos. 3,993,995 (Kaplan et al, 
1976) and 3,524,058 (Robertson et al, 1970). 
SUMMARY OF THE INVENTION 
The present invention provides a respiration monitor, and X-ray trigger 
apparatus which is not subject to erroneous identification of 
non-breathing related artifacts as the respiration waveform, or subject to 
false indication of apnea. 
A DC coupled, frequency insensitive preamp provides a true, non-phase 
distorted representation of the sensed analog respiration signal. An 
undistorted representation of the ECG signal (diagnostic quality) is also 
provided. An adaptive DC offset correction whereby a predetermined 
percentage (e.g., 90%) of the DC offset is subtracted from the analog 
signal is performed. The adaptive DC offset correction, in effect, 
improves the information to DC offset level (percent modulation) by a 
factor of 10, and reduces the voltage range of the analog signal to a 
point where digitalization becomes practicable. A digital representation 
of the respiration cycle is thus provided. The ECG signal is also 
digitized. 
The digital signals are analyzed to determine respiration extrema, 
respiration rate and heart rate. Apnea episodes are indicated and/or other 
alarms actuated when the detected respiratory rate decreases by a 
predetermined percentage (e.g., 50%), in successive breaths, accompanied 
by a predetermined number (e.g., 4) of successive periods of heart rate 
deceleration; if the respiration period exceeds a predetermined duration; 
if the respiratory rate exceeds operator set maximum and minimum threshold 
values; if the heart period exceeds a predetermined duration; if the heart 
rate exceeds operator set maximum or minimum threshold values; or if the 
detected respiratory rate approximately equals the detected heart rate for 
more than a predetermined number of periods. 
Artifact rejection is provided by the DC coupling of the system and by 
amplitude, slew rate and time domain discrimination based upon known 
physical parameters of the respiration waveform and cardiac components. 
Generation of an alarm if detected respiration and heart rates are equal 
provides a fail-safe against erroneous identification of the cardiac 
component as the respiration waveform. In addition, an indication of the 
relative quality of electrode contact and thus of the data is provided.

DETAILED DESCRIPTION OF A PREFERRED EXEMPLARY EMBODIMENT 
Referring now to FIG. 1 there is shown a block diagram of a respiration 
monitor in accordance with the present invention. A respiration sensor 198 
provides output voltage directly indicative of the respiration cycle of a 
patient. Respiration sensor 198 may be any device which provides a signal, 
directly indicative of respiration. However, in the preferred embodiment, 
respiration sensor 198 is an impedance pneumograph. An impedance 
pneumograph sensor 198 is particularly advantageous in that measurement of 
the ECG can be effected through the same electrodes. 
In impedance pneumography, a low level, high frequency carrier signal is 
injected into the thorax. A voltage is developed which is indicative of 
the transthoracic impedance of the subject. The transthoracic impedance 
changes as a function of the respiration cycle. The injected carrier may 
be of a wide range of frequencies and amplitudes, limited primarily in 
that the signals should be below neural response levels. Sinusoidal, 
squarewave or triangle waveform signals are suitable. 
In the presently preferred exemplary embodiment a 30 kHz squarewave pulse 
train of 10 .mu.sec. width pulses is used. The carrier provides a constant 
current level on the order of 200 .mu.a. A squarewave carrier modulation 
waveform is utilized to facilitate use of a phase synchronous demodulator, 
as will be explained. 
As noted above, the transthoracic impedance varies in accordance with the 
respiration cycle. The transthoracic impedance includes a resistive 
component generally associated with pulmonary blood flow. The 
transthoracic impedance also includes a capacitive component generally 
associated with the changing capacitance of the thorax as the volumed mass 
ratio of the thorax changes during the respiration cycle. 
The carrier signal is injected and the resultant voltage is sensed through 
electrodes electrically coupled to the patient's thorax. A bipolar (two 
electrode) configuration wherein the same set of electrodes both inject 
the carrier and sense the resultant voltage, or a tetrapolar (four 
electrode) configuration, using separate sets of electrodes for excitation 
and voltage sensing may be utilized. 
The signal provided by impedance pneumograph comprises an AC information 
(respiration waveform) signal with a large DC offset. The resting 
impedance of the thorax (measured on a bipolar electrode configuration) is 
generally on the order of 400-1000 ohms at frequencies between 50 Hz and 1 
MHz. The AC varying component of the impedance signal representing the 
respiratory cycle is generally between 0.2 ohms to 5 ohms. Thus, there is 
only on the order of 0.05% modulation. Breathing rates vary from on the 
order of six breaths a minute to on the order of 180 breaths a minute. 
Thus, the AC information component of the respiration signal is of 
relatively low frequency ranging from nearly DC. 
In addtion, the AC component of transthoracic impedance, includes a 
"cardiac component" artifact, attributable to the operation of the heart, 
rather then respiration. The cardiac component has a frequency in 
accordance with heart rate and normally has an amplitude equating to and 
from approximately 0.02 ohms to 0.2 ohms. 
The respiration signal from sensor 198 is applied to a preamplifier 200. 
Preamplifier 200 provides a high input impedance to the respirator sensor, 
and, in effect, DC couples the respiration sensor to the remainder of the 
monitor system. Preamp 200 operates both on the respiration impedance 
signal and the ECG signal. Preamp 200 demodulates the information signal 
from the carrier to develop a raw analog respiration waveform signal. 
Similarly, the ECG waveform signal is derived from the sensor output. 
Historically, respiration monitors have been AC coupled to the respiration 
sensor. A large capacitance (RC time constant on the order of 8 seconds) 
has been coupled across the monitor inputs to filter out DC offsets in the 
input signal. However, intelligence in the signal is often at frequencies 
as low a 0.1 Hz. Where the signal is AC coupled into the monitor, the 
frequency response of the monitor is typically -3 db at 0.5 Hz. 
Accordingly, the information components at low frequencies are greatly 
attenuated. When respiration rate decreases to, for example, 30 breaths 
per minute, the frequency response of conventional respiration monitors is 
such that only approximately 70% of the signal amplitude is sensed by the 
monitor. Thus, a 0.3 ohm transthoracic impedance would appear to be only 
0.21 ohms to the detector. Similarly, when the respiratory rate falls to, 
for example, 15 breaths per minute, the response has fallen on the order 
of -6db. Accordingly, the apparent transthoracic impedance registered by 
the monitor is only on the order of half the actual value. Thus, as the 
subject's respiratory rate decreases, the magnitude of the transthoracic 
impedance registered by the monitor steadily decreases. In addition, the 
frequency response of the prior art monitor tends to cause erroneous 
interpretation of the cardiac component of transthoracic impedance as the 
respiratory waveform. During apnea P.sub.aCO.sbsb.2 increases, 
(hypercapnea) causing a steady deceleration of heart rate from an initial 
value on the order of 140 beats per minute (approximate upper frequency -6 
db point) to rates as low as 60 breaths per minute. By virtue of the 
frequency response of the system, the decreasing heart rate effects a 
two-fold increase in the registered magnitude of the cardiac component of 
transthoracic impedance. This phenomena is complemented by an actual 
increase of cardiac component amplitude resulting from the increased 
stroke volume of the heart at lower rates. The cardiac component thus 
drops into the frequency range normally associated with respiration and 
increases in amplitude to values corresponding as much as 0.2 ohms. In 
prior art AC coupled systems, a low pass filter (-3 db at 1.5 Hz) is 
generally used to discriminate against the cardiac component. However, 
during hypercapnic periods frequency discrimination becomes ineffectual 
and the cardiac component becomes indistinguishable from the respiratory 
component of transthoracic impedance. Accordingly, apneic episodes are 
often masked by erroneous detection of the cardiac component as the 
respiratory waveform. In fact, FDA regulations require all presently 
available impedance pneumograph sensor apnea monitors to include a warning 
to the effect that there is a possibility that the apparatus will monitor 
heart rate during severe apneic episodes. In addition, the respiration 
waveform is both time and amplitude distorted by the passband of the prior 
art systems. The typical passband of conventional respiration monitors 
peak at frequencies corresponding to 60 breaths per minute (having a 
nominal voltage equivalent to 1 ohm of transthoracic impedance). However, 
a wide range of respiration rates are exhibited by, for example, an infant 
(6 breaths per minute to 180 breaths per minute). It has been found that 
in most commercially available systems the electrical signal after 
coupling lags the actual respiration cycle by on the order of 30%-45% at 
30 breaths per minute and undergoes a full phase reversal at on the order 
of 120 breaths per minute. DC coupling the respiration waveform to the 
monitor avoids distortion and phase shift in the respiration signal. 
Preamp 200 also provides over voltage protection to guard the system 
against damage in the subject while respiration is being monitored. In the 
preferred exemplary embodiment, preamp 200 also includes provisions for 
common mode rejection, a preamplifier shield and a shield driver to 
eliminate radiated power line frequency signal problems. Preamp 200 will 
be described more fully in conjunction with FIG. 2. 
The respiration waveform produced by preamplifier 200 is applied to a 
respiration analog processing circuit 300. Respiration analog processing 
circuit 300, provides waveshaping and DC offset compensation for the 
signal. Adaptive subtraction of approximately 90% of the DC component in 
the respiration waveform signal is performed. It has been calculated from 
the human physiology that as much as 91.6% of the DC component of the 
respiration waveform could be subtracted without entering into the range 
of values affected by the intelligence (AC) component. However, the DC 
base line very often drifts. Accordingly, an adaptive substraction 
technique must be utilized. Respiration analog processing circuit 300 will 
be described in more detail in conjunction with FIG. 3. 
The processed analog signal is applied to arithmetic processing circuitry 
400. Arithmetic processing circuitry 400 includes a 16-bit 
analog-to-digital A/D converter, digital peak an slope detector 404 and a 
time domain rate discriminator 500. Peak detector 404 and time domain rate 
discriminator 500 operate as artifact rejection circuitry. The theory of 
operation is based upon knowledge of human physiology, to wit, limitations 
on the amount and rate of change of transthoracic impedance during the 
respiratory cycle and the relative timing of respiratory events. Digital 
peak and slope detector 404 provides signals indicative of inspiration and 
expiration states during the respiratory cycle. These signals may be 
utilized to provide audio/visula indicia of the subject's respiration. 
Changeovers between inspiratory and expiratory periods are indicative of 
maxima and minima in the respiratory cycle. 
Digital peak and slope detector 404 is also, in effect, a digital filter, 
providing amplitude discrimination against changes in impedance not of 
sufficient magnitude to be associated with respiration. In particular, the 
resolution of the monitor with respect to respiration extrema is chosen to 
be 0.3 ohms. Impedance changes of less than 0.3 ohms are effectively 
filtered out by digital peak and slope detector 404 and a divide-by-2 
counter. Assuming the electrodes to be properly positioned, the digital 
filtering successfully discriminates against cardiac components of ths 
signal. The cardiac component is generally on the order of 0.1 ohm. In 
addition, the maximum rate of change of thoracic impedance (or 
respiration) has been empirically determined for various populations of 
subjects. Accordingly, digital peak and slope detector 404 also rejects 
impedance changes exceeding the maximum possible change during a sampling 
period. For an infant, the maximum respiration rate is 180 breaths per 
minute (equating to a respiration rate of 3 Hz). A 25 msec. sample period 
is chosen in accordance with Shannon sampling therorum. Accordingly, if 
the signal changes at a slew rate greater than 1 ohm per 25 second sample, 
the signal can not relate to physical respiration and accordingly, the 
system is reset. 
Time domain discriminator circuit 500 is responsive to the minima and 
maxima signals from peak detector 404 and operates to inhibit the monitor 
with respect to signals not occurring within the known time constraints of 
human physiology. Digital peak and slope detector 404 and time domain 
discriminator circuitry 500 will hereinafter be more fully described in 
conjunction with FIGS. 4 and 5. 
A time discriminated extrema signal is applied to a respiration tachometry 
circuit 600. Respiration tachometry circuit 600 computes the actual 
respiration rate of the subject. The respiration tachometry circuit 600 is 
used to selectively activate an audio visual alarm if the respiration rate 
drops below or exceeds preset limits, and further, drives displays 
indicating the average or instantaneous respiration rate. Respiration 
tachometry circuit 600 also provides artifact rejection by discriminating 
against respiration period counts differing from the just previously 
measured period by more than a predetermined amount. The predetermined 
amount can be fixed, or can be adaptive, based upon predetermined 
percentage of the previous period. 
Digital peak and slope detector 404, time domain discriminator 500 an 
respiration tachometry circuit 600 all provide input signals to an X-ray 
trigger circuit 700. X-ray trigger circuit 700 will be described in 
conjunction with FIG. 7. X-ray trigger circuit 700 triggers the X-ray 
machine slightly before a selected respiration extrema. X-ray trigger 
circuit 700 provides for triggering of an X-ray machine only at a 
predetermined point in the respiratory cycle during a valid respiration 
measurement. It is first determined that a steady state of respiration 
exists, and no alarm conditions are present. In addition, the changes in 
successive respiratory periods must be within acceptable limits. 
Physiologically, the human respiratory system can not change rate by more 
than a predetermined percentage, e.g., .+-.20%, from the preceding period. 
It has been determined that the percentage can be equated to a 
predetermined number of breaths per minute for various population groups, 
e.g., infants, children, adults, etc. For infants, it has been found that 
20 breaths per minute is an effective upper limit on rate changes between 
breaths. Further, the amplitude variations must be within prescribed 
limits, i.e., within a prescribed percentage of the last measured 
respiration cycle or an equivalent maximum or minimum change. Lastly, the 
X-ray technician must maintain the trigger circuit in an armed state by 
maintaining a switch in a depressed condition. If all of these criteria 
are met, trigger circuit 700 detects the transthoracic impedance at a 
point in the respiration cycle corresponding to the desired specie of 
respiration extrema, adjusts the measured impedance by a predetermined 
account to provide lead time for the X-ray machine, then triggers the 
X-ray machine when the adjusted value of impedance is next sensed in the 
respiratory cycle. X-ray trigger circuit 700 thus provides for accurate 
triggering of the X-ray machine at a desired level of respiration. The 
accuracy of triggering, as a practical matter, reduces the total exposure 
necessary to provide good X-ray pictures. X-ray trigger circuit 700 is 
also designed to prevent the X-ray machine from taking more than one 
exposure without rearming. In addition, if a valid trigger point is not 
detected in five successive breaths, the arm cycle will terminate and 
rearming (i.e., again depressing the rearm button) necessitated. The ECG 
signal developed by preamp 200 is applied through suitable analog 
processing circuitry 800 to a cardiotachometer 900. 
Heart rate tachometry circuit 900 is similar to respiration tachometer 600 
and provides a signal indicative of heart rate. The heart rate signals 
utilized to drive an audio visual alarm when the heart rate exceeds or 
drops below respective maximum and minimum threshold values. A display of 
the heart rate is also provided. 
Respiration monitor system 100 generates apnea alarms when any of the 
following circumstances are detected: (a) when an apneic period (time 
period between respective inspirations) exceeds a preset time period; (b) 
when a respiration period exceeds the preceding respiration period by a 
predetermined factor and a predetermined number of successive decreases 
are detected in the heart beat period (physiological apnea); and (c) if 
the respiration rate and heart rate detected are equal for more than one 
sample (indicative of erroneous detection of the cardiac component as 
respiratory waveform). 
As previously mentioned, alarms indicative of improper respiration and 
heart rates are also provided. In addition, an apnea alarm is also 
generated if a predetermined number of successive heart beats are at rates 
less than 100 (for infants). 
Monitoring system 100 includes a number of provisions for preventing false 
alarms, without denigrating the sensitivity of the system to physical 
apnea. Amplitude, slew rate and time domain discrimination based upon 
empirical determinations of maximum and minimum values of physical 
parameters are utilized. For example, in the analog processing circuit 
300, any voltage change corresponding to a change greater than, for 
example, .+-.20 ohms from the instantaneous DC base line is rejected as 
not being associated with normal respiration, i.e., an artifact and the 
system reset. Similarly, digital peak and slope detector 404, in the 
capacity of a digital filter, rejects any amplitude changes corresponding 
to less than, e.g., 0.15 ohms. Similarly, slew rate discrimination is 
provided based on the premise that it is physically impossible for the 
respiration waveform to increase at greater than a predetermined maximum 
rate. Accordingly, changes in transthoracic impedance greater than, for 
example 1 ohm during the course of a sample period (e.g., 25 msec.) are 
rejected. Further, time domain discrimination relating to the relative 
timing of particular events during the course of the respiratory cycle is 
provided by time domain rate discriminator 500. Similarly, an upper limit 
on respiratory rate has also been empirically determined. Accordingly, an 
upper threshold for valid respiratory rate measurements is established in 
respiratory tachometry circuit 600. 
In addition, provisions are made to prevent erroneous interpretation of the 
cardiac component as the respiration waveform. Respiration monitor 100 is 
DC coupled, that is, presents a perfectly flat frequency response with 
respect to low frequency respiration signals. Accordingly, the amplitude 
drop off evidenced in the prior art systems at low respiration rates does 
not occur in monitor 100. Further, the amplitude discrimination of digital 
peak and slope detector 404 effectively prevents mistaken identification 
of the cardiac component as the respiration waveform. Assuming the 
apparatus to be properly used and electrodes properly disposed on the 
subject, the cardiac component never reaches the 0.3 ohm sensitivity level 
of monitor 100 and, erroneous identification of the cardiac component as 
the respiratory waveform can not occur. 
In addition, as a fail-safe, and apnea alarm is energized if the respiraton 
and heart rates (as sensed by monitor 100) are equal for more than a 
predetermined number of sequential sample periods. The heart rate and 
respiration rate are physically asynchronous, and accordingly, in terms of 
statistical analysis should not track. Accordingly, if the sensed 
respiration rate and the cardiac rate appear to be equal, it is assumed 
that the cardiac component has bee mistakenly identified as the 
respiration waveform. Such mistaken identity can occur not withstanding 
the amplitude discrimination for example, in the event that electrodes are 
improperly placed, resulting in an abnormally high cardiac component. 
It should also be appreciated that DC coupling of monitor 100 to the 
respiratory sensor provides a faithful, substantially unphase-shifted 
respiration waveform, to facilitate accurate measurement of inspiration to 
expiration (I/E) ratios. The inspiration to expiration (I/E) ratio is a 
parameter desirable for adjustment of artificial ventilators for use on 
infants. Where the ventilator is adjusted to provide 
inspiration/expiration ratio not in accordance with the natural I/E ratio 
of the infant, the infant tends to fight or resist the ventilator. Monitor 
100 determines an average I/E ratio for the subject used only with valid 
data and exclusing all statistically significant false information. 
In addition, monitor 100 can readily be adapted to control the ventilator. 
The parameters of the subject's respiration cycle are determined during 
normal breathing periods by monitor 100. When the breathing significantly 
varies from those parameters, i.e., breathing becomes labored, the 
respirator is set up in a feedback loop with monitor 100 to rematch the 
normal respiration of the subject. 
The respirator/respiration monitor control accommodates the problem of 
weaning the infant from a respiratory. The respirator can be programmed to 
turn off the respirator at intermittent periods, reactivating the 
respirator at the end of a predetermined period or in response to a 
particular physiological effect, such as decreasing heart rate. 
Referring now to FIG. 2, a more detailed description of a preferred 
preamplifier 200 will be provided. As previously noted in the preferred 
embodiment, respiration sensor 198 comprises an impedance pneumograph. A 
constant current pulse train is applied across the varying thoracic 
impedance to generate a voltage which varies in accordance with the 
transthoracic impedance. A set of electrodes 202 are disposed on the 
subject in a conventional bipolar or tetrapolar configuration. In a 
bipolar configuration, a set of electrodes, e.g., E1 and E2, are disposed 
on the mid-auxiliary lines over the sixth intercostal space. Electrode E1 
is connected to one connector of a constant current source 204. Constant 
current source 204 is inductively coupled to an oscillator 206. Oscillator 
206, in effect, on/off modulates current source 204 to provide a 300 kHz 
squarewave modulated signal of constant current pulses on the order of 200 
.mu.a in amplitude and 125 .mu.sec. pulse width. An exemplary current 
source 204 is shown in schematic form in FIG. 2a. The second electrode, 
e.g., E2 disposed on the opposite side of the thorax is connected to the 
other conductor of constant current source 204. A neon lamp 208 is coupled 
across the output lines of current source 204 to provide for over-voltage 
protection in the event of application of defibrillation signals to the 
subject. A third electrode, E3 is disposed on the subject, preferably 
equidistant to electrodes, to provide for detection of common mode 
signals. 
Electrodes E1, E2 and EC are all coupled to suitable over-voltage 
protection circuitry 210. Over-voltage protection circuitry 210 is 
primarily to protect against damage to the monitor in the event of 
defibrillation procedures on the subject. The over-protection circuitry, 
in effect, couples electrodes E1 and E2 and the common reference electrode 
EC to the respective input terminals of a high input impedance 
differential amplifier 212. Suitable over-protection circuitry 210 and 
high input impedance amplifier 212 are shown in schematic form in FIG. 2b. 
The output signal of differential amplifier 212, provided across output 
terminals 212a and 212b, comprises an on/off modulated constant current 
carrier (125 sec. pulses and duty cycle 30Hz) amplitude modulated by the 
changing thoracic impedance. The output terminals 212 and 212a are coupled 
across the primary of an isolating transformer 214 (1:1). 
The secondary of transformer 214 is coupled to a conventional phase 
synchronous demodulator 216. Demodulator 216 is also receptive of a signal 
indicative of the excitation carrier waveform from blocking oscillator 
206. Phase synchronous demodulator 216, in effect operating as an homodyne 
demodulator, provides a demodulated output signal indicative of the 
respiratory waveform without any substantial phase shifting. 
A suitable synchronous demodulator is shown in schematic form in FIG. 2c. 
In effect, demodulator 216 operates on the principle of turning an FET on 
and off in synchronism with the pulses of oscillator 206, storing a charge 
on a capacitor indicative of the peak valve of the squarewave. The voltage 
across the capacitor is representative of the demodulated waveform without 
any significant phase shift. 
The demodulated signal, generated across terminals 216b and 216c are 
applied to an amplifier 218 and low pass filter 220 to provide a 
differential analog signal to respiration analog processing circuit 300. 
Suitable amplifier 218 and low pass filter 220 are shown in schematic form 
in FIG. 2d. Amplifier 218 is adjusted to provide a sensitivity of 10 mV 
per ohm. Low pass filter 220 filters out high frequency transients in the 
demodulated waveform. At the frequencies associated with respiration, low 
pass filter 220 injects essentially no phase shift into the signal. The 
amplified, filter, demodulated signal is provided at output terminals 220a 
and 220b for application to respiration analog processing circuit 300. 
Preamplifier 200 includes provisions for compensating for common mode 
signals (60 cycle noise). In effect, the common mode signals are measured 
and negatively fed back to the input of the preamp. The output terminals 
of amplifier 212 are coupled to a balanced differential amplifier 222, 
which provides an output signal equivalent to the common mode signal. The 
common mode signal is then inverted by an inverting amplifier 224 and 
capacitively coupled to reference electrode EC. Negative feedback is thus 
established to cancel common mode signals. 
To further eliminate any outside electromagnetic interference in monitor 
100, monitor 100 is encased in a three layer driven shield schematically 
illustrated and denoted in FIG. 2 as 196. Shield 196 suitably comprises an 
internal shield of 0.015 hydrogen annealed 2680 permalloy. The common mode 
voltage is applied to the internal shield as a driving signal. An 
intershield of 1/8" teflon or equivalent high di-electric material and an 
outer layer of fine mesh copper screen tied to earth are also utilized. It 
has been found that driven shielding 196 substantially eliminates 
electromagnetic interference with monitor 100. 
ECG signals are also derived from the output of amplifier 212. The ECG is a 
biopotential present on the subject's body. The ECG signals generally 
comprises voltage having frequency components ranging from DC to on the 
order of 100 Hz. The ECG signals are blocked from respiration demodulator 
216 by 0.1 .mu.f blocking capacitor 226, interjected between one lead of 
the primary of transformer 214 and output terminal 212a of differential 
amplifier 212. Capacitor 226 and the primary of transformer 214 cooperate 
as a high pass filter, blocking frequencies above the order of 100 cycles. 
The output terminals of amplifier 212 are connected to a low pass filter 
228. Low pass filter 228 suitably comprises a pie-section filter 
comprising a 10 mH inductor and 0.1 .mu.f capacitor. The output of low 
pass filter 228 is taken across the capacitor. Low pass filter 228, in 
effect, filters out the carrier components of the signal and generates a 
slowly timed variant DC voltage having a frequency content ranging from 
zero to on the order of 100 Hz. 
It is desirable that the patient be electrically isolated from power ground 
to prevent electrical shock and for greater common mode rejection at power 
line frequencies. However, the output of low pass filter 228 is a very 
slowly varying signal not generally compatible with passage through a 
transformer. Accordingly, the ECG signal is, in effect, chopped by a phase 
synchronous modulator 230, driven by a signal derived from blocking 
oscillator 206. Thus, the output of modulator 230 is essentially the low 
frequency (DC) output signal of low pass filter 228 chopped at a rate 
synchronous with block oscillator 206. A suitable phase synchronous 
modulator is shown in schematic form in FIG. 2e. The chopped signal is 
applied across transformer 232 to a phase synchronous demodulator 234, 
similar to phase synchronous demodulator 216. 
Ringing between modulator 230 and demodulator 234 is prevented by utilizing 
field effect transistors having different threshold voltages, e.g., 2N5654 
FET'S having a pinch-off of -8 volts are utilized in modulator 230 and on 
FET-type 2N5664 (having a pinch-off voltage of -6 volts) is utilized in 
demodulator 234. By appropriate choice of relative pinch-off voltages 
demodulation is effected during relatively noise-free (transient) periods 
in the output signal of modulator 230. 
The output of the ECG demodulator 234 is applied to a suitable amplifier 
236 and a low pass filter 238. Amplifiers 236 and 238 are shown in 
schematic form in FIG. 2f. The frequency response of filter 238 is 
essentially flat with respect to frequencies from DC to 100 Hz, the -3 db 
point being at 100 Hz. When the frequency is above 500 Hz, the response 
rolls off at 80 db per decade. The frequency response may be generally 
categorized as a Butterworth flat frequency response. While some phase 
shifting is effected by the filter, the shifting does not significantly 
effect the geometry of the ECG waveform. It should be appreciated that the 
common mode rejection scheme, driven shield and high isolation impedance 
described above, provide for generation of an ECG waveform without 
requiring 60 cycle filtering. Accordingly, the output signal from low pass 
filter 230 is of diagnostic quality. 
The respiration waveform provided across outputs 220a and 220b of low pass 
filter 220 is applied to respiration analog processing circuit 300. 
Respiration analog processing circuit 300 provides signal conditioning, 
adaptive subtraction of a predetermined percentage (e.g., 90%) of the DC 
offset level and indicia of the relative quality of electrode connection. 
The signals from preamp LPF 220 are applied to a variable gain 
differential amplifier 302. The single ended output of amplifier 302, 
provided at output terminal 302a, is connected to an inverting amplifier 
304 and to one leg of a summing circuit 306, as will be explained. 
Variable gain differential amplifier 302 standardizes and linearizes the 
sensitivity response of monitor 100. Sensitivity is adjusted to exactly 
10mV per ohm, with 1 V equivalent to 100 ohms and 10 V corresponding to 
1000 ohms transthoracic impedance. Variable gain differential amplifier 
302 and inverting amplifier 304 are together shown in schematic form in 
FIG. 3c. Linearity and temperature stability of response are ensured 
through use of military grade LH2101AD/833B monolithic operational 
amplifiers. 
The output (raw analog) of inverting amplifier 304 (provided at 304a) is 
analyzed with respect to impedance level to determine the relative quality 
of the electrode connections to the subject. The raw analog signal at 
terminal 304a is applied to a plurality of comparators 308-311. The other 
inputs to comparators 308-311 are receptive of respective voltages 
corresponding to predetermined impedances. For example, comparator 308 has 
applied to 10 V reference corresponding to a measured transthoracic 
impedance of 1000 ohms. Comparators 309 and 310 operate with reference 
voltages corresponding to, for example, 700 ohms and 400 ohms 
transthoracic impedance, respectively. Comparator 311 operates with 
reference voltage corresponding to just under 100 ohms (e.g., 0.98 V), the 
low end of acceptable transthoracic impedance values. The respective 
outputs a-d of comparators 308-311 are applied to a conventional decoder 
312, which selectively drives indicator lamps corresponding to quantitized 
electrode connection ratings. Under normal operating conditions the 
minimum thoracic impedance encountered is on the order of 100 ohms. 
Accordingly, where the respiration waveform indicates a thoracic impedance 
of less than 100 ohms, a malfunction exists. For example, the electrodes 
may be shorted together. Similarly, where transthoracic impedances over 
1000 ohms are encountered, an improper electrode connection to the subject 
or a system malfunction exists. A truth table for decoder 312 is provided 
in Table 1. 
TABLE I 
______________________________________ 
Z&gt;1000 Z&gt;700 Z&gt;400 Z&gt;100 
A B C C Indicator 
______________________________________ 
0 0 0 0 Low Z, bad electrode 
connection 
0 0 0 1 Excellent electrode 
connection 
0 0 1 1 Good electrode 
connection 
0 1 1 1 Fair electrode 
connection 
1 1 1 1 High Z, bad electrode 
connection 
______________________________________ 
It should be appreciated that the indication of the relative quality of 
electrode contact is also an indication of the validity of the data. In 
addition, the outputs of comparator 308 (indicative of transthoracic 
impedance greater than 1000) and the inverted output of comparator 311 
(the inverted output being indicative of transthoracic impedances of less 
than 100) are applied to an OR gate 314. The output of OR gate 314 may be 
termed the Z out-of-limits signal. 
The Z out-of-limits signal is applied to a master clear one-shot 316, which 
generates a negative going pulse ouput signal to reset the system in the 
event of an impedance measurement out of physiological limits. Master 
one-shot 316 is also triggered upon powering up of the system by suitable 
circuitry 318 to initiate the system. Power-on clear circuitry 318 
suitably comprises a RC timing circuit which provides a high level signal 
through OR gate 314 to master clear one-shot 316 until the capacitor in 
the circuit charges after a predetermined period whereupon the output 
assumes a low level. Master clear 316a synchronizes the respective system 
components upon powering up and resets the respective contents upon an 
indication of an out-of-limits impedance. 
As noted above, respiration analog processing circuitry 300 compensates for 
DC offset. In effect, 90% of the DC voltage (a percentage determined not 
to impinge upon any information signal) is subtracted from the raw analog 
signal. The raw respiratory analog signal provided at terminal 304a is 
applied to a voltage divider 316. Voltage divider 316 is adjusted to 
provide an output signal equal to 0.9 (90%) of the applied input signal. 
The 90% signal from voltage divider 316 is applied through a low pass 
filter 318 to a voltage follower buffer amplifier 320. Low pass filter 318 
filters out any high frequency noise transients. Voltage follower 320 
operates as a buffer and provides a greater current capability. Exemplary 
circuitry for voltage divider 316, low pass filter 318 and buffer 
amplifier 320 is shown in schematic form in FIG. 3d. Buffer amplifier 320 
is of linear response with no DC offsets. To this end, in the preferred 
embodiment, a military grade LH2101AD/8838 operational amplifier is 
utilized. 
The output of buffer amplifier 320 provided at output terminal 320a is 
applied to a single slope A/D converter 322. A/D converter 322 generates a 
pulse width modulated output signal indicative of the amplitude of the 90% 
signal. In effect, A/D converter 322 compares a digitally generated linear 
ramp signal to the 90% signal, and generates a pulse having a first 
transition at the beginning of the sample period and a second transition 
when the ramp voltage equals the sample. The sample rate is chosen to be 
once every 10 seconds. The digital slope generator is driven by a clock 
signal developed from a free running oscillator and divider 324. An 
exemplary A/D converter 322 is shown in FIG. 3e. 
Clock 324 and single slope A/D converter 322 together operate as a pulse 
width modulator 326 providing a pulse width modulated signal and a signal 
indicative of the end of the 10 second sampling period. 
The pulse width modulated signal, provided at output 322b, the end of 
sample signal, provided at terminal 322c and the Z clock signal, provided 
at terminal 322a, are applied to suitable count and control logic 328. 
Logic 328 generates, at respective output terminals 328a, 328b, 328c and 
328d, respective signals designated Zcount, Zlatch, Zload, and Zreset. 
These signals are illustrated in FIG. 3f. Briefly, the clock signals from 
terminal 322a are gated during the duration of the pulse width modulated 
signal provided at terminal 322b, to produce the Zcount signal at terminal 
328a. Zlatch, Zload, and Zreset are sequential signals generated in timed 
relation with the operational cycle of respiration analog processing 
circuit 300. Suitable logic for 328 is shown in FIG. 3g. 
The Z counts provided at terminal 328a increment a binary counter 330. The 
Z counts are also applied as clock signals to an up/down counter 332. 
Up/down counter 332 and binary counter 330 are connected in parallel 
through a register of latches 334, as will hereinafter be explained. 
Latches 334 are used to store a count representative of the DC base line 
for the respiration waveform. Latches 334 are selectively loaded with the 
binary count when the 90% count of a particular sample differs from the 
previous base line count by a predetermined amount. More particularly, the 
parallel outputs of up/down counter 332 are applied as one set of inputs 
to a magnitude comparator 336. The other set of inputs to comparator 336 
are hardwired to correspond to a predetermined value in accordance with 
the maximum excursion of information. For example, the thoracic impedance 
doesn't generally change by more than .+-.5 ohms. Accordingly, a value 
equivalent to an impedance of 10 ohms can be utilized as the threshold 
value. In the preferred embodiment, a value equivalent to 30 ohms is 
utilized to prevent resetting of the system by spurious noise. At the end 
of the pulse width modulated signal a count has been accumulated in binary 
counter 330 indicative of the 90% analog value. Latches 334 are initially 
loaded with the contents of binary counter 330 in response to the next 
Zlatch signal from logic 328 in coincidence with an arithmetic enable 
signal generated, as will be explained, by arithmetic processing circuitry 
400 in predetermined timed relation with the master clear one-shot pulse 
from one-shot 316. Thereafter, in response to the Zload signal, up/down 
counter 332 is loaded with the contents of latches 334. As previously 
noted, the contents of latches 334 represent the 90% level at sample time. 
Binary counter 330 is thereafter reset in response to the reset signal 
from logic 328. During the next sampling period, binary counter 330 is 
incremented and up/down counter 332 is decremented in response to the Z 
counts. At the end of the accumulation period, the count in up/down 
counter 332 is indicative of the magnitude of the difference between the 
previous count, i.e., the 90% base line, and the instantaneous sample. 
Accordingly, if the difference exceeds the predetermined threshold, the 
instantaneous sample is then loaded in response to the Zlatch control 
signal into latches 334 to update the DC base line, and a pulse is applied 
through OR gate 314 to trigger master clear one-shot 316 and reset the 
system. If the difference between the sample and base line count does not 
exceed the predetermined threshold, it is assumed that the difference is 
attributable to information and the 90% DC base line in latches 334 
maintained. Thus, the contents of latches 334 are indicative of the 
sampled 90% level. 
The contents of latches 334 are also reconverted into an analog signal by 
an 8-bit D/A converter 338. The analog signal is applied through a 
suitable scaling amplifier 340 and active low pass filter 342 to the 
second leg of summing junction 306. Scaling amplifier 340 restores the 
output signal from D/A converter 338 to the 90% analog level. Scaling 
amplifier 340 is suitably an operational amplifier having a high input 
impedance. Filter 342 blocks any switching transients in the output signal 
of D/A converter 338. An exemplary scaling amplifier 340 and active low 
pass filter 342 are shown in schematic form in FIG. 3h. 
Summing circuit 306, adds the 90% DC base line and the raw analog signal 
from terminal 302a. Summing junction 306 is formed by precision metal film 
resistors 207 and 309. Summing junction 306a is coupled to a voltage 
follower buffer amplifier 344. A suitable buffer amplifier 344 is shown in 
FIG. 3i. 
The output of voltage follower 344 is applied to an inverting amplifier 
346, having a gain of -20. Inverting amplifier 346 provides an output 
signal of proper polarity. Amplifier 346 is connected to a low pass filter 
348 to further reject any transients in the signal, without providing any 
appreciable phase shifting of intelligence. A suitable inverting amplifier 
346 and low pass filter 348 are shown in schematic form in FIG. 3j. 
The output of low pass filter 348a represents the raw analog signal minus 
90% of the sampled DC signal level. More particularly, the signal 
represents the instantaneous signal (ac.sub.i +dc.sub.i) minus 90% of the 
signal (Ac+Dc) at sample time divided by two, times the gain of amplifier 
346. Assuming the AC component to be negligible with respect to the DC 
component at sample time, it is found that the signal is approximately 
equal to ten times the AC component (information) plus the DC component. 
Thus, the intelligence to DC offset ratio is increased by a factor of 10. 
The output of analog processing circuit 300 provided at terminal 348a will 
hereinafter be referred to as the 1% analog signal. 
It is significant that the entire circuit is DC coupled and that very 
little phase shifting is effected on the signal. The frequency response of 
all of the low pass filters were chosen to have virtually no effect on the 
AC intelligence signals. 
The output of analog processing circuitry 300 provided at output terminal 
348a is applied to arithmetic processing circuitry 400. Arithmetic 
processing circuitry 400 develops signals indicative of the occurrence of 
respiration extrema, and performs, amplitude, slew rate and time domain 
discrimination on the data to reject artifacts. 
The 1% analog signal is applied to a 16-bit single slope A/D converter 
operating, in effect, as a pulse width modulator. A suitable 16-bit A/D 
converter is shown in schematic form in FIG. 4c. 
As previously noted, arithmetic processing circuitry 400 includes a digital 
peak and slope detector 404. Digital peak and slope detector 404 includes 
a 5MHz oscillator 405, cooperating with a 24-bit binary counter 406. The 
parallel outputs of counter 406 are applied to suitable timing decoder 
logic 408. Timing decoder logic 408 generates respective sequential timing 
signals ZCEP1, ZCEP2, and ZCEP3 at output terminals 408a, 408b and 408c, 
respectively. 
A shift register 410 is clocked by signals from the last bit of counter 
406. Shift register 410 is utilized to ensure that the processing 
circuitry reaches steady state before calculations are performed and 
generates the previously mentioned arithmetic enable signal at terminal 
410a. 
Digital peak and slope detector 404 is initialized by the master clear 
pulse from master clear one-shot 316. The master clear one-shot pulse is 
applied as a clear signal to counter 406 and to shift register 410. As 
previously described, the master clear one-shot is generated in response 
to powering up of the system, and out-of-limits impedance (indicative of 
improper electrode connections to the subject) or a change in the DC 
offset level. Shift register 410 effects a delay from initialization to 
ensure that all counters in the system have reached steady state, i.e., 
the counters have had time to be purged of false initial noise counts. 
A 40 Hz sampling rate is established by the signal from bit 16 of counter 
406. Bit 16 is coupled to input 402a of 16-bit A/D converter (pulse width 
modulator) 402. More specifically, bit 16 has a period of 25 msec., thus 
establishing 12.5 msec. convert period and a 12.5 msec. calculation period 
within the overall sampling period. At the beginning of the 12.5 msec. 
convert period, a linear ramp is generated in A/D converter 402. When the 
linear ramp reaches the value of the 1% analog signal at terminal 348a, 
the convert period is terminated, thus providing a pulse width modulated 
output signal at terminal 402b of 16 bit A/D converter 402. The pulse 
width modulated signal is inverted and applied to one input terminal of a 
two input NOR gate 436. The other input of NOR gate 436 is coupled to the 
5 MHz clock signal. Thus, the pulse width modulated signal gates the clock 
signals to provide a number of pulses (hereinafter referred to as a 
Zcount) indicative of the 1% analog signal. The Z count is provided at 
terminal 436a. 
The Z count and the respective timing signals from timing logic 408 are 
applied to suitable control logic 430 and steering gates 432 and 434, 
which alternately apply the Z counts and the timing signals from timing 
logic 403 to respective banks A and B of computation apparatus. Suitable 
control logic 430 and steering gates 432 and 434 are shown in schematic 
form in FIGS. 4d, 4e and 4f, respectively. 
Computation apparatus A comprises a binary counter 412a, coupled in 
parallel with an up/down counter 414. Up/down counter 414a is in turn 
coupled in parallel with a plurality of comparators 416a, 422a, and 427a. 
Comparators 420a, 422a and 427a have coupled to the second set of inputs 
thereof hardwired preload values 418a, 424a and 428a, respectively. The 
A&gt;B outputs of comparators 416a and 422a and the A&lt;B output of comparator 
428, are respectively applied to input terminals of a 4-input AND gate 
420a, and three 3-input AND gates 426a and 429a. The second and third 
input terminals of AND gates 420a, 426a and 429a have applied thereto the 
ZCEP1 signal and the ZCTDNA signal generated at terminal 432a of steering 
gates 432, respectively. The fourth input terminal of AND gate 420.beta. 
is responsive to the A&lt;B output of conparator 428a. 
The second computation apparatus (computation apparatus .beta.) comprises a 
binary counter 412.beta., up/down counter 414.beta., comparators 
416.beta., 422.beta. and 427.beta. and hardwired preload 418.beta., 
424.beta. and 428.beta., 4-input AND gate 420.beta. and 3-input AND gates 
426.beta. and 429.beta.. Comparation apparatus .beta. is essentially 
identical to computation apparatus A, with the exception that AND gates 
420.beta., 426.beta. and 429.beta. are responsive to the CNTDN.beta. 
signal generated at terminal 434b of steering gates 434, and in the 
control signals to counter 412.beta. and up/down counter 414.beta.. 
Counters 412a and 412b are each 16-bit counters to accommodate numbers up 
to 65,000 (corresponding to a 10 V maximum 1% analog signal at sensitivity 
10 ohms per volt, and A/D ramp timing such that 650 counts are generated 
per 100 mV). 
Computation apparatus A and computation apparatus .beta. are utilized to 
provide amplitude discrimination against signals not associated with 
respiration, and in particular, the cardiac component. Two banks of 
calculation apparatus are utilized to provide for comparison of each 
sample to the just preceding sample. Control logic 430 and steering gates 
432 and 434 apply the respective control signals to calculation apparatus 
A and calculation apparatus .beta. to effect the following operation. The 
counters are initially reset in response to the arithmetic enable signal 
from terminal 410a. Assuming computation apparatus A to be initially 
addressed, in response to the ZCEP2 control signal, the contents of binary 
counter 412a will be parallel loaded into up/down counter 414a. Binary 
counter 412b is reset by ZCEP3 from terminal 402c. During the next 
sampling period, the Z count is applied to increment binary counter 412b 
and to decrement up/down counter 414a. Thus, at the end of the conversion 
period: up/down counter 414a contains a count indicative of the absolute 
difference between the successive samples. The final state of the sign bit 
indicates whether the absolute difference is a positive or negative 
number; and binary counter 412b contains a count indicative of the last 
taken sample. 
The difference count in up/down counter 414a is compared to the hardwired 
preload value of 418a. Preload 418a is chosen to be equivalent to 0.15 
ohms, one-half the smallest amplitude change typically encountered in the 
physical respiration waveform. Comparator 416a provides a signal 
indicative of whether the Z count sampled has changed sufficiently from 
the last Z count to be considered valid. 
As noted above, the A greater than B output (A&gt;B) of comparator 416a is 
applied to a one input of 4-input AND gate 420a. The other inputs of AND 
gate 420a are the ZCEP1a signals from timing logic 408 and terminal 408a, 
and countdown A (CTDNA) control signal from steering gates 432 
(originating in control logic 430), and the A&lt;B output signal from 
comparator 427a, as will be explained. The output terminal 421a of AND 
gate 420a is coupled to control logic 430. The output of AND gate 420a 
will hereinafter be referred to as F(A)r. Generation of the F(A)r signal 
causes control logic 430 to toggle, in effect, reversing the roles of 
computation apparatus A and computation apparatus B. 
If the difference count in counter 422 is not greater than the preload 
threshold count corresponding to 0.15 ohms, the relative roles of 
computation apparatus A and computation apparatus B are maintained as they 
were. The previous contents of binary counter 412a are again loaded into 
up/down counter 414a for comparison against the next successive sample (Z 
count) while the next sample is accumulated in binary counter 412b in 
computation apparatus B. Thus, one of computation apparatus A or 
computation apparatus B accumulate the Z count sample while the other 
subtracts the Z count from the previously accumulated sample and compares 
the difference to the predetermined count corresponding to one-half the 
desired sensitivity of the system. Any change between samples less than 
the preload threshold is thus, in effect, filtered out. Upon detection of 
a valid sample, calculation apparatus A and calculation apparatus B change 
roles, under the control of control logic 430 and steering gates 432 and 
434. Computation apparatus A and computation apparatus B are to all 
effects and purposes identical, and, for the purposes of simplicity, the 
corresponding elements for the respective calculation banks will 
hereinafter be referred to generically by their numerical designation. For 
example, it will be understood that when comparator 416 is referred to as 
providing an A greater than B output to AND gate 420, that reference is 
made to analogous connections between comparator 416a and GATES 420a and 
between comparator 416b and AND gate 420b. 
Differing sensitivities can be provided by comparing the difference between 
successive samples against different sensitivity preloads. For example, 
preloads 424 can be chosen to establish a sensitivity of 0.01 ohms. 
Accordingly, by gating in AND gate 426, the respective A greater than B 
(A&gt;B) outputs of comparators 422 with ZCEP1 from terminal 408a of timing 
logic 408 and the countdown A control signal (CTDNA) from terminal 432b of 
steering gates 434 (or countdown B from terminal 434b), the cardiac 
component can be monitored to digitally develop a heart rate signal in a 
manner analogous to the analysis of the respiration waveform, as will be 
explained. It should be appreciated that such digital generation of a 
heart rate signal completely eliminates 60 cycle interference. With 
respect to monitor 100, this is a moot point in view of the shield driver 
previously described. However, in some instances certain cardiac diseases 
or anomalies greatly distort the ECG waveform; conventional ECG monitoring 
systems reject the waveforms as artifacts. In some instances, a digital 
generation of an heart rate signal could be desirable. 
A still further comparison can be made against a preloaded threshold value 
equivalent to the maximum amount allowable change. That is, a comparison 
can be made to a count corresponding to the maximum slew rate 
physiologically possible in the human respiratory cycle. To this end, the 
A less than B (A&lt;B) output of comparators 427 are used as inputs to AND 
gates 429 and 420. Assuming a maximum breath rate of 180 breaths a minute, 
the maximum change in transthoracic impedance during the sample period is 
less than 1 ohm. Accordingly, by setting the preload 428 threshold value 
at 1.5 ohms and inhibiting generation of the valid sample signals, F(A)r 
or F(B)r when the difference count exceeds the threshold by AND gates 420, 
high slew rate artifacts are rejected. A sample is therefore accepted as 
valid only in the event that it differed from the preceding valid sample 
by an amount exceeding one-half the sensitivity level of the system as 
established by preload 418, but also within acceptable maximum limits as 
determined by preload 428. Thus, both amplitude and slew rate 
discrimination are provided against artifacts. 
Digital peak detector 404 provides an indication of the occurrence of 
respiration extrema. Signals indicative of the sign (zero crossover) of 
the difference counts in up/down counters 414a and 414b are applied to the 
respective inputs of a 2-input NAND gate 442. It should be appreciated 
that the sign of the difference between succesive samples is indicative of 
the direction of slope of the respiration waveform between those sampling 
instances. Accordingly, the zero crossovers in up/down counters 414 and 
422 are indicative of the occurrence of a respiration extrema. When either 
of up/down counters 414 transition from down counting to up counting 
(crossover zero), a low level sign signal is generated to NAND gate 442, 
forcing the output thereof to go high. The output of NAND gate 442 will 
hereinafter be referred to as the up/down flag signal. 
The up/down flag signal and the valid count signals F(A)r (421a) and F(B)r 
(421b) are applied to a slope detector 445. The up/down flag signal is 
applied to the data input of a D flip-flop 444 and to the clear input of a 
flip-flop 446 operating as a divide-by-2 circuit. Flip-flop 446 is clocked 
in response to transitions between the A and B computation apparatus. The 
clock signal to flip-flop 446 is provided by a 2-input NOR gate 448 having 
the respective inputs thereof connected to terminals 421a and 421b, 
respectively. Divide-by-2 flip-flop 446 adjusts the effective sensitivity 
of the system to 0.3 ohms (hence the setting of preload 418 at 0.15 ohms). 
The output of flip-flop 446 is utilized to clock flip-flop 444. The Q and 
Q outputs of flip-flop 444 are used to drive an inspiration indicator and 
expiration indicator, respectively. 
So long as the samples are increasing in value, i.e., the slope is 
positive, a zero crossover (sign change) will occur. Similarly, if no zero 
cross-over occurs, it is indicative that the successive samples are 
decreasing in value, i.e., the slope is negative. The output of D 
flip-flop 444 reflects the up/down flag signal and is thus indicative of 
the direction of slope of the respiration waveform. The Q input provides 
an indication of positive slope (terminal 444a) and the Q output provides 
an indication of negative going slope (terminal 444b). 
The positive slope and negative slope output signals provide at terminals 
444a and 444b of digital peak detector 404 are applied to time 
discrimination circuitry 500. Time discrimination circuitry 500 
discriminates with respect to artifacts occurring outside the 
physiological time constraints of the respiration cycle. For example, 
maximum and minimum time periods between the relative events in the 
respiration cycle have been empirically established, and signals not 
falling within the physiological limits are rejected as artifacts. 
Referring now to FIG. 5, time domain discriminator 500 will be described 
in more detail. 
Time domain discriminator 500 comprises a plurality of one-shots: two 800 
nsec. one-shots, a 111 msec. one-shot 558, and a 222 msec. one-shot 562. 
All of the one-shots are enabled by the calculate enable signal provided 
at terminal 430c of control logic 430 (connections not shown). The pulse 
widths of the one-shot outputs are chosen in accordance with the 
particular subject population on which the apparatus is to be used. The 
pulse widths herein described, relate to infants. The pulse widths would 
differ for different subject populations such as older children or adults. 
The maximum physiologically possible breath rate can be established for 
the population (180 breaths per minute in infants). Accordingly, 
successive inspiratory or expiratory peaks can not reoccur within a period 
of on the order of 333 msec. Further, inspiratory to expiratory ratio are 
on the order of 1-to-2, that is, expiration time is twice the inspiration 
time. It therefore follows that a minimum period of on the order of 111 
msec. occurs between an inspiratory peak and an expiratory peak. 
Similarly, a minimum period on the order of 222 msec. will pass between 
the occurrence of an expiratory peak (maxima) and inspiratory peak 
(minima). The relative time relationships are illustrated in graphic form 
in FIG. 5a. 
When the slope of the respiratory waveform (1% analog signal) changes from 
negative to positive, indicative of a minima in the waveform (expiratory 
peak) flip-flop 444 changes state, generating a high output at positive 
slope terminal 444a. The positive going transition causes one-shot 554 to 
trigger (assuming one-shot 554 is not inhibited by one-shot 560, as will 
be explained). The 0.8 .mu.sec. pulse generated by one-shot 554 at 
terminal 554a will hereinafter be referred to as the low-up pulse. It must 
be appreciated that in view of the 0.15 ohm preload in 418 (establishing a 
0.3 ohm sensitivity) the occurrence of a low-up pulse corresponds to a 
point on the respiratory waveform within 0.3 ohms of the minima. Taking 
the worst case, the figure 0.3 arises as follows. Referring to the 1% 
analog respiration waveform of FIG. 5a, assume that the minima corresponds 
to a thoracic impedance of 100 ohms. Assume a first sample S1 is taken and 
a count corresponding to 98.85 ohms is accumulated in binary counter 512, 
and that a negative slope signal is provided by flip-flop 444. Assume the 
next successive sample S2 occurs at the peak (100 ohms). The difference is 
thus 0.15 ohms. When compared to the preload sensitivity threshold value 
0.15, an A greater than B signal is not generated, and sample S2 ignored. 
The 98.85 count is thus retained in counter 412 and loaded into up.down 
counter 414 for differencing with the next successive sample S3. Assuming 
sample S3 to correspond to a transthoracic impedance of 98.85 ohms, that 
sample S2 will be filtered out by the amplitude discrimination. The next 
sample, however, S4 (here 98.6 ohms) will cause a zero crossover in 
up/down counter 414b causing generation of the up/down flag and a positive 
transition in the positive slope signal at terminal 444a. Thus, a low up 
pulse is generated at terminal 554a by one-shot 554. The low up pulse 
triggers 111 msec. one-shot 558 and 333 msec. one-shot 560. The output of 
one-shot 560 is fed back to one-shot 554 as an inhibit signal to establish 
the minimum relative timing between the occurrences of low-up pulses 
(respiration minima). Any apparent negative to positive slope transition 
occurring within 333 msec. of the last previous negative to positive 
transition, must be due to an artifact, and accordingly, one-shot 554 is 
inhibited with respect to the second transition. One-shot 558 establishes 
the minimum relative time between the negative to positive transition and 
positive to negative transition. 
Similarly, 0.8 .mu.sec. one-shot 556 is triggered by positive going 
transitions in the negative slope by one-shot 556, at terminal 556a, is 
indicative of a positive to negative slope transition in the respiration 
waveform and will hereinafter be referred to as the high down pulse. 
One-shot 556 is similarly inhibited by 333 msec. one-shot 556 to 
discriminate against apparent positive to negative slope transitions 
occurring too soon after a previous negative to positive transition. 
One-shot 562, triggered by the high down pulse generates a pulse 
indicative of the minimum time between positive and negative and negative 
to positive slope transitions in the waveform. Further, since the time 
periods represented by one-shots 558 and 562 are mutually exclusive, the 
outputs of the respective one-shots are utilized to inhibit each other. 
As previously noted, the one-shots of time domain discriminator 500 are 
enabled by the calculate enable signal from terminal 430c of control logic 
430. The calculate enable signal is generated by a shift register within 
logic 430 and operates to insure that the system is in steady state before 
any alarm signals or X-ray trigger signals are generated. 
The 333 msec. low-up to low-up pulse generated by one-shot 560 in response 
to the low-up pulse is gated to respiration tachometry circuitry 600. 
Respiration tachometry circuit 600, in effect, measures the period between 
the leading edges of successive low-up to low-up pulses, to determine the 
respiration rate. Both the instantaneous rate, (the rate during each 
successive interval), and an indication of the rate averaged over a 
predetermined number of events are provided. 
Referring now to FIG. 6, and in particular FIG. 6a, a more detailed 
description of respiration tachometry circuit 600 will be provided. The 
gated pulse indicative of the occurrence of low-up (terminal 564a) is 
applied to a waveshaper 602. A suitable waveshaper 602 is shown in 
schematic form in FIG. 6e. Waveshaper 602, in effect, standardizes the 
input signal at a proper logic voltage level. Waveshaper 602 generates at 
an output terminal 602a an "event" signal, synchronous with the leading 
edge of the one-shot pulse generated at terminal 564a. In addition, almost 
simultaneously with the generation of the event signal, waveshaper 602 
generates a narrow pulse Rbeat at terminal 602b (and Rbeat at terminal 
602b). Rbeat synchronizes the rest of the respiration tachometry circuit 
600, generated in response to the first master clock signal occurring 
after the "event". The clock signals are generated by a master clock 604 
having a time constant of 578 .mu.sec. The timing of the output signals of 
the respective elements of waveshaper 602 are shown in FIG. 6f. 
The master clock signal is also applied to a timing driver chain 608 which 
generates at respective output signals 608a, 608b and 608c, respective 480 
pulse per second, 30 pulse per second, and 15 pulse per second signals. 
The Rbeat output provided at terminal 602c of waveshaper 602 is applied to 
the clear input of a shift register 610. Shift register 610 is clocked by 
the 15 pulse per second signal from timing chain 608, and establishes a 
logically true signal Rconvert for a 66 msec. period following Rbeat, 
hereinafter referred to as the convert period. The Rconvert signal is 
generated at terminal 610a and the Rconvert signal generated at terminal 
610b. The Rconvert signal is applied to suitable timing logic 612, 
together with the 480, 30 and 15 pulse per second signals from timing 
chain 608, to generate respective control signals RCEP1 at terminal 612a, 
RCEP2 at terminal 612b, RCEP3 at terminal 612c, Rload at terminal 612d and 
RCEP1a at terminal 612e. Suitable timing logic 612 is shown in FIG. 6g. 
The relative time relationship of the respective signals are shown in 
diagrammatic form in FIG. 6h. 
The Rconvert signal (610b) and the 480 pps signal (608a) are applied to a 
period to rate converter 620. In effect, converter 620 accumulates a count 
at the 480 pps rate indicative of the period between respective beat 
signals in a counter 614. The accumulated count is loaded into a down 
counter 616. Counter 616 is thereafter repetitively counted down at the 
clock frequency during the convert period. Each time down counter 616 
counts down to zero, a most significant borrow bit is generated and passed 
through a delay 618 to count generator logic 619. The delayed most 
significant bit is fed back to down counter 616 to cause the counter to 
reload with the period count accumulated in counter 614. The number of 
times down counter 616 counts down (i.e., the number of msb bits 
generated) during the convert period is indicative of the instantaneous 
respiration rate. 
In practice, the 480 pulse per second signal is gated through a NAND gate 
613 to clock counter 614 only during the convert period. Accordingly, 
binary counter 614 is preloaded with a count corresponding to the known 
(66 msec.) duration of the convert period, i.e., 38 counts. Binary counter 
614 is reset by the RCEP1a control signal generated at terminal 612a of 
timing logic 612. 
Suitable count generator logic 619 is shown in FIG. 6i. Count generator 
logic 619, in effect, gates the most significant borrow bit from up/down 
counter 616 during the convert period to generate an RPDO count. The 
number of RPDO counts is indicative of the instantaneous respiration rate 
in breaths per minute. 
It should be appreciated that other varieties of period to rate converters 
can be utilized. However, period to rate converter 620 is particularly 
advantageous in that 7-segment LED displays can be directly driven without 
the necessity of look-up tables, as will be explained. 
Counter 614 also generates an R overflow signal in the case of an anomaly, 
that is, if the Rconvert period (indicative of the respiration period) 
extends beyond predetermined bounds. Counter 614 can accumulate a count up 
to 2048 (corresponding to 16 seconds). If the respiration period exceeds 
16 seconds, an R overflow signal is generated at terminal 614a. Generation 
of R overflow resets the counters to zero, activates an apnea alarm and 
clears display enable logic 622 (as will be described). 
It should be appreciated that it requires the occurrence of two events 
(low-up pulses) to provide a valid computation of respiration rate. 
Accordingly, the display is blanked until a valid calculation is attained. 
The display blanking function is performed by display enable logic 622. 
Display enable logic 622 is responsive to the event signal from waveshaper 
602 and to the R overflow signal provided by counter 614 at terminal 614a. 
Display enable logic 622 generates respective control signals RDE1, RDE2, 
RDE3 at terminals 622a, 622b and 622c. Display enable logic 622 is 
essentially a shift register with data input tied high, clocked by the 
event pulses and reset by R overflow (614a) or master clear. 
Artifact rejection, that is discrimination against spurious counts is also 
provided in respiration tachometry circuit 600 by artifact rejection logic 
664, shown in FIG. 6b. It is physiologically impossible for the breath 
rate of a subject to increase by more than a predetermined amount during 
the course of a sampling period. The predetermined amount may be an 
empirically determined maximum for the particular population group of the 
subject, e.g., plus 20 beats per minute for an infant, or may be 
adaptively determined on a percentage of previous rate basis. Accordingly, 
when a rate count exceeds the previous rate count by more than the 
permissible amount, that rate count is deemed to be a spurious count and 
is rejected. Further, provisions must be taken not to incorporate the 
spurious count into the computation of average rates. However, it must be 
appreciated that the maximum change limit pertains only to increasing in 
respiratory rate; apnea may occur almost instantaneously. 
In addition, the possibility of a spurious initial count must be 
considered. Accordingly, if a predetermined number of successive rate 
counts exceed the base count by more than the predetermined limit, the 
base count is assumed to be an error and the later rate counts used as a 
new base for computation. 
With reference now to FIG. 6b, artifact rejection logic 664, will now be 
described. The RPDO counts from period to rate converter 620 (terminal 
619a) are applied to one input of a 2-input NAND gate 659. NAND gate 659 
gates the RPDO counts to clock an 8-bit binary counter 660. Counter 660 is 
reset by the RCEP3 signal generated at terminal 612c of timing logic 612. 
The parallel outputs of counter 660 are in turn applied to the inputs of 
an up/down counter 672. The outputs of up/down counter 672 are applied to 
one set of inputs of a magnitude comparator 680. The other set of inputs 
of magnitude comparator 680 cooperate with a hardwired preload 681 
representative of a predetermined maximum limit on the amount of change 
physiologically possible between respective events (breaths). It should be 
appreciated that rather than a fixed maximum, an adaptive system whereby a 
difference is compared to a predetermined percentage of the preceding 
sample may be utilized. 
After the system has warmed up, i.e., 2 events have occurred, NAND gate 659 
gates the RPDO counts to counter 660. Thus, at the end of each convert 
period, counter 660 has accumulated a count indicative of the 
instantaneous respiration rate. 
Latches 662 establish the base respiration count against which subsequent 
counts are compared. The base respiration count is the last occurring 
valid (within the system constraints) rate count sampled. As noted above, 
the rate count is deemed a valid sample if it does not differ from the 
preceding base rate count (contents of latches 662) by more than the 
predetermined limit established by preload 681. The rate count is 
similarly deemed to be valid if the sample differs from the preceding base 
count by more than the predetermined limit, but is a decreasing rate, or 
if a predetermined number (e.g., 5) of preceding rate samples have been 
deemed invalid as out-of-limits (indicating that perhaps the previous base 
count was spurious). 
The accumulation of a valid count in counter 660, is significed by the 
generation of the RCEP1a-1 signal (valid count) generated at terminal 666a 
of logic 666. The accumulated count in counter 660 is selectively latched 
in latches 662 to update the base count in response to the RCEP1a-1 
signal. The RCEP1a-1 signal is generated by a 2-input OR gate 668. One 
input of OR gate 668 is receptive of the output signals from a 2-input AND 
gate 669. The inputs of AND gate 669 are responsive to the RDE3 signal 
derived from terminal 662c of display enable logic 662 and the RCEP1a 
signal generated at terminal 612e of timing logic 612. 
The other input of OR gate 668 is responsive to the output signal of a 
2-input AND gate 670. One input of AND gate 670 is responsive to the 
RCEP1a signal at terminal 612e. The other input terminal of AND gate 670 
is responsive to the output of a 3-input OR gate 671. The respective 
inputs to OR gate 671 are responsive to the output of a 2-input AND gate 
694 operating upon the A greater than B (A&gt;B) from comparator 670 and the 
countdown signal RCNTDN (as will be described); to the A less than B (A&lt;B) 
signal generated by magnitude comparator 680 at terminal 680b; and to a 
"force" signal generated by the QE (bit 5) output of a shift register 690 
(FIG. 6c). Shift register 690 is indexed by the RCEP1 signal of timing 
logic 612 and is reset by RCEP1a-1. The QB (second bit) output is provided 
to X-ray trigger circuit 700 (arming logic 704), such that one spurious 
count will interrupt the arm cycle. 
During the first three periods of operation (during initial start-up or 
after a reset), the contents of counter 660 is transferred to the output 
of latches 662 to establish an initial rate base count. For the first 
three periods, signal RDE3 is high. A high RDE3 forces AND gate 669 high. 
Accordingly, OR gate 668 generates a high level RCEP1a-1 signal, causing 
the content of counter 660 to be latched over into latches 662. 
The difference between the instantaneous respiration rate sample and the 
base rate is established by up/down counter 672. After each convert period 
control signal Rload is generated by timing logic 612. The Rload signal is 
applied as a load command to up/down counter 672, causing the contents of 
latches 662 to be loaded into up/down counter 672. The up/down control of 
up/down counter 672 is effectd through a JK flip-flop 676, having J input 
tied high and K input tied to ground. Flip-flop 676 is clocked by the 
max/min output (occurring at zero count) of counter 672 and is cleared by 
the Rload signal generated by timing logic 612. The Q output of flip-flop 
676 is tied to the down control of up/down counter 672. The Q output of 
flip-flop 676 also provides the count down signal to AND gate 694. Thus, 
at the same time counter 672 is loaded with the contents of latches 662, 
flip-flop 676 is cleared to place up/down counter 672 in a down count 
mode. The next set of RPDO counts count up/down counter 672 down. At the 
end of the convert period, up/down counter thus contains a count equal to 
the difference between the stored count and the instantaneous period 
count. It should be appreciated that in the event up/down counter 672 
reaches zero, flip-flop 676 is clocked, causing the Q output to go high 
and Q to go low. Up/down counter 672 will thus enter an up count mode. In 
addition, at the end of the convert period the instantaneous period count 
has been accumulated in counter 660. 
The difference count in up/down counter 672 is compared to the preloaded 
threshold by magnitude comparator 680 and an A greater than B (A&gt;B) or A 
less than B (A&lt;B) signal generated accordingly. If A is less than B (A&lt;B) 
the present count accumulated in counter 660 is latched by latches 662 at 
the RCEP1a time. That is, if A is less than B, OR gate 671 provides a high 
level output signal at terminal 671a to enable AND gate 670 with respect 
to the RCEP1a latch control signal. Accordingly, OR gate 668 generates the 
RCEP1a-1 output signal at terminal 666a, causing latches 662 to latch the 
accumulated count in counter 660. 
It is desired, however, that the maximum limit on rate change be applied 
only to increases in respiratory rate. Accordingly, where the second rate 
is less than the previous rate count, the lesser value is latched into 
latches 662 irrespective of the absolute difference in the rates. A 
decrease in rate is signified by retention of the down count mode of 
operation by down counter 672 at the end of the convert period, and, thus 
by a count down signal (RCNTDN) generated at the Q output of flip-flop 
676. Accordingly, when the difference between respective rate counts 
exceeds the maximum increase threshold (A greater than B) and a high level 
down count signal is provided by flip-flop 676 Q, AND gate 694 generates a 
high level output signal, causing OR gate 671 to enable AND gate 670 with 
respect to the RCEP1a signal. Upon the next occurrence of the RCEP1a 
signal the RCEP1a-1 signal will therefore be generated at terminal 666a, 
causing latches 662 to latch the contents of counter 660 (the second rate 
count). 
As noted above, a possibility exists that an erroneous base count is 
established in latches 662. An erroneous base count arises primarily 
through the initial establishing of the base count in latches 662 during 
the first three samples of operation after reset. Accordingly, if a 
predetermined number of sequential samples are deemed out of limits, it is 
assumed that the present base rate is in error and the subsequent rate 
latched as the new base rate. This is accomplished by shift register 690. 
Shift register 690 is indexed in response to each RCEP1 signal, indicative 
of each sample taken, and is reset by the RCEP1a-1 signal. Thus, if five 
samples are taken, that is, shift register 690 is indexed by five 
successive RCEP1a signals, without an intervening reset RCEP1a-1 pulse 
(indicative of a valid sample), the QE output of shift register 690 goes 
high to generate a "force" signal to OR gate 671. The high level output of 
OR gate 671 enables AND gate 670 with respect to the next occurring RCEP1a 
pulse. Accordingly, the next RCEP1a pulse is passed through AND gate 670 
and OR gate 668 to generate the RCEP1a-1 signal causing the last rate 
count accumulated in counter 660 to be latched in latches 662 as the new 
base rate and resets shift register 690. 
As previously noted, both an average respiration rate and the instantaneous 
respiration rate are selectively displayed. The selective averaging is 
provided by select averaging logic 624 shown in FIG. 6c. Selective 
averaging logic 624 comprises a divide-by-5 circuit 626 operative on the 
RPDO counts from period to rate converter 620. The RPDO counts and the 
output of the divide-by-5 circuit 626 are applied to multiplexed inputs of 
a multiplexer 628 to provide a selected RPDOS count at the output terminal 
of the 628a of the multiplexer 628. Similarly, the Rconvert signal (610b) 
from shift register 610 clocks a divide-by-t counter 630 to provide an RD5 
output signal. The RD5 output signal is applied to one input of each of 
two 2-input NAND gates 632 and 634. The other input of NAND gate 632 is 
coupled to the output of (670a) of 2-input AND gate 670. The output (670a) 
of AND gate 670 will hereinafter be referred to as the valid RCEP1a-1 
signal. The valid RCEP1a-1 signal is applied together with the output 
signal (valid RCEP1a-1D5) of NAND gate 632 to two multiplexed inputs of 
multiplexer 628 to provide a RCEP1a-1S output signal at terminal 628b of 
the multiplexer. The second input of NAND gate 634 is receptive to the 
RCEP3 signal generated at terminal 612c of timing logic 612. The output of 
NAND gate 634 (RCEP3D5) and the RCEP3 signal are applied to two 
multiplexed inputs of multiplexer 628 to provide a RCEP3S signal at 
terminal 628c of the multiplexer. Counter 630 is reset by the output of a 
3-input NAND gate 638. NAND gate 638 is responsive to the A greater than B 
signal generated by magnitude comparator 680, the R count up signal 
generated by flip-flop 676 at terminal 676a and the RCEP1a signal 
generated at timing logic terminal 612c. 
The average is taken over five respiration periods. The five respiration 
period averaging interval is established by, in effect, dividing the 
Rconvert signal generated at terminal 610b of shift register 610, by five. 
It should be recalled that the RPDO counts are generated during the 
convert time, the number of counts generated during that time being 
indicative of the rate. Accordingly, the Rconvert signal from terminal 
610b is applied to a divide-by-5 counter 630. Divider 630 generates a high 
level output signal RD5 during the fifth Rconvert period. The RD5 signal 
is applied to NAND gates 632 and 634 to gate one-in-five valid RCEP1a-1 
pulses to generate the signal RCEP1a-1D5 and one-in-five RCEP3 pulses to 
generate the signal RCEP3D5. Simultaneously, divide-by-5 counter 626 is 
generating output pulses in response to one-in-five RPDO counts. 
Accordingly, during the averaging mode, multiplexer 628 provides at its 
outputs 628a, 628b and a pulse train indicative of one-fifth (1/5) of the 
RPDO counts, every fifth valid RCEP1aD5 count (i.e., every fifth output 
pulse from AND gate 670) and every fifth RCEP3 reset pulse. The time 
relationship of the respective signals relating to selective averaging 
logic 624 are shown in FIG. 6j. 
In the event of a spurious count (A&gt;B and R counting up) the operation of 
selective averaging logic 624 is inhibited, by resetting divide-by-5 
counter 630. Application of the reset pulse to divide-by-5 counter 630 
causes counter 630 to assume all 1's generating a high RD5 signal until 
the next negative transition of Rconvert. The high level RD5 signal resets 
counter 626 and causes generation of RCEP3D5 upon the next occurring RCEP3 
pulse, to effect reinitialization of various components of an alarm 
circuit 640. 
Alarm circuit 640, shown in FIG. 6d, comprises an inverting BCD counter 
642. BCD counter 642 suitably comprises three 74LS90 integrated circuits. 
BCD counter 642 is clocked by the RPDOS signal provided at multiplexer 
output terminal 628a, and is reset by the RCEP3S signal provided at 
multiplexer output terminal 628c. The parallel outputs of BCD counter 642 
are coupled through diodes to a low limit BCD thumb-wheel switches 644. 
The parallel outputs of BCD counter 642 are also applied through diodes to 
high limit BCD thumb-wheel switches 645. Thumb-wheel switches 644 and 645 
generate high level output signals when the count in BCD counter 642 
equals the respective setting on the thumb-wheel switches. 
The output of low limit thumb-wheel switches 644 is applied to the clock 
input of a D flip-flop 646. The D input of flip-flop 646 is tied to 
ground, and the preset input responsive to RCEP3S, generated from the 
RCEP3S signal at multiplexer terminal 628c. Flip-flop 646 is thus 
initially set, providing a high Q output signal, and is clocked low when 
the BCD count equals the low limit set in thumb-wheel switches 644. The Q 
output of flip-flop 646 is applied to the J input of a JK flop 647. The K 
input of flip-flop 647 is responsive to the inverted Q output signal Q of 
flip-flop 646. JK flip-flop 647 is clocked by the RCEP1a-1S signal 
generated at multiplexer output 628b, and is preset by the R overflow 
signal generated by counter 614 at terminal 614a. The Q output of 
flip-flop 647 is applied as a control signal to an audio visual alarm 648. 
The output of high limit thumb-wheel switches 645 is applied to the clock 
input of a D-type flip-flop 649. D-type flip-flop 649 is connected in a 
configuration identical to flip-flop 646 with D input tied to ground and 
preset responsive to the RCEP3S signal derived from the output of 
multiplexer terminal 628c. The Q output of flip-flop 649 is applied 
through an inverter 650 to the J input of a JK flip-flop 651, (so shown 
for emphasis in distinguishing from flip-flops 646 and 647). JK flip-flop 
651 is clocked by the RCEP1a-1S signal provided at multiplexer output 
628b. The K input of flip-flop 651 is receptive of the Q output signal 
from flip-flop 649. 
In operation, when the count in BCD counter 642 equals the low limit set on 
thumb-wheel switches 644, flip-flop 646 is clocked, causing the Q output 
thereof to change from high to low. Accordingly, a low J-input and high K 
input are provided to flip-flop 647. Thus, upon the next subsequent 
RCEP1a-1S time, flip-flop 647 will generate a low Q output and a high Q 
output. Thus, if the count in BCD counter 642 exceeds the low limit set on 
thumb-wheel switches 644 during the appropriate accumulation time, no 
alarm is generated. 
However, if at the end of the accumulation time, (convert or RD5) 
associated with RPDOS, the low limit set on thumb-wheel switches 644 has 
not been reached, no clock signal is applied to flip-flop 646, and 
accordingly, at the RCEP1a-1S time, the inputs to the J and K terminals of 
flip-flop 647 are values one and zero. Accordingly, when clocked by the 
RCEP1a-1S pulse and the Q output goes high and an audio visual alarm is 
initiated. 
Audio visual alarm 652 is activated if the high limit threshold set in 
thumb-wheel switches 645 is reached by the count in BCD counter 642. If 
the count is not reached, no alarm is sounded. When the content of BCD 
counter reaches the high limit threshold, thumb-wheel switches 645 clock 
flip-flop 649 causing the Q output thereof to go low. Accordingly, at the 
RCEP1a-1S time, a one is applied to the J input and a zero applied to the 
K input of flip-flop 651. When clocked by the RCEP1a-1S pulse, flip-flop 
651 therefore generates a high level Q output signal to initiate an alarm. 
In normal steady state operation, (no high or low alarm conditions present) 
the RCEP1a-1D5, RCEP3D5 and RPDOD5 signals are provided at multiplexer 
output terminals 628b, 628c and 628a. The alarm function therefore, is 
operating on the basis of a five breath averaged interval. However, in the 
event of a high or low limit alarm condition occurring, it is desirable to 
operate thereafter in response to the instantaneous respiration rate until 
the average rate returns to normal limits. To this end, the select input 
to multiplexer 628 operates in response to the Q output of a JK flip-flop 
653. The J input of flip-flop 653 is responsive to the output of a 2-input 
NOR gate 654. The inputs to NOR gate 654 are coupled to the Q outputs of 
flip-flops 647 and 651. The K input of flip-flop 653 is responsive to the 
output of a 2-input NAND gate 655. The inputs of NAND gate 655 are RD5 
(derived from the RD5 signal provided at output terminal 630a of selective 
averaging logic 624) and the output of NOR gate 654. Flip-flop 653 is 
clocked by the RCEP2 signal derived from the RCEP2 signal provided at 
timing logic output terminal 612b. When no alarm condition exists, the Q 
outputs of flip-flops 647 and 651 are both low. Accordingly, the output of 
NOR gate 654 is high providing a high input to the K input terminal of 
flip-flop 653. The RD5 signal is high and NAND gate 655 provides a low 
input at the K input terminal of flip-flop 653. Accordingly, the RCEP2 
signal clocks the Q output of flip-flop 653 high causing multiplexer 628 
to select the appropriate inputs for the averaging mode operation. 
However, if an alarm condition occurs, at least one of the Q outputs of 
flip-flop 647 and 651 will assume a high value. The output of NOR gate 654 
will therefore go low providing a low J input to flip-flop 653 and forcing 
NAND gate 655 to provide a low K input to flip-flop 653. Thus, the Q 
output of flip-flop 653 goes low at RCEP2 time, causing multiplexer 628 to 
provide the instantaneous signals. RCEP1a-1, RCEP3 and RPDO at terminals 
628b, 628c and 628a. Alarm circuit 640 thereafter operates under the 
control of the instantaneous rate signals. 
When the alarm conditions no longer exist, it is desirable that the 
operation of alarm circuit 640 return to the control of the average 
signals. However, control cannot be returned to the average signals until 
the beginning of an averaging period as indicated by the RD5 signal. 
Accordingly, when a no-alarm condition is reassumed, and RD5 goes low, 
high level signals are generated to the J and K inputs of flip-flop 653. 
Thus, when clocked by the RCEP2 signal, flip-flop 653 toggles causing 
multiplexer 628 to select the average signals until the next averaged 
alarm condition. 
Provisions are made for display at either the instantaneous or averaged 
respiration rate. A second multiplexer 656 receives the same input signals 
(RCEP1a-1D5, valid RCEP1a, RCEP3D5, RCEP3, RPDO5, RPDO) as multiplexer 
628. The respective output terminals are applied to conventional driver 
circuits which inturn provide an input to a conventional BCD to seven 
segment converter 657. BCD to seven segment converter 657 drives display 
621. 
Multiplexer 651 is controlled by suitable MUX inhibit logic 657. MUX 
inhibit logic 657 is responsive to the RD5 signal generated by selective 
averaging logic 624 at terminal 630a, the RCEP2 signal generated at timing 
logic output terminal 612b, and to a pushbutton average select switch 658. 
Suitable MUX inhibit logic is shown in FIG. 6k. MUX inhibit logic 657 
operates to ensure that transfer from display of instantaneous to average 
rates occurs at the beginning of an averaging period. The high down and 
low up pulses provided at terminals 554a and 556a are also applied to 
X-ray trigger circuit 700 shown in FIG. 7, to effect maxima and minima 
triggering, respectively. 
As previously noted, it is often necessary that an X-ray picture be taken 
at a respiration extrema. Automatic triggering of the X-ray machine at the 
proper instant of time, particularly with respect to a subject incapable 
of cooperation such as an unconscious adult or an infant is therefore 
desirable. In addition, a finite time period is required after the 
application of a trigger signal to the X-ray machine for the actual 
charging and discharging of the high voltage on the anode which emits the 
X-rays. The period is a function if, inter alia, exposure time and 
intensity of the X-ray. Accordingly, it is desirable that the trigger 
voltage be generated at an instant leading the particular respiration 
extrema at which triggering is to take place. X-ray triggering circuit 
700, in effect, stores a Z count corresponding to the sample occurring 
just after a chosen one of the high down or low up pulses are generated, 
and adds or subtracts therefrom a count equivalent to a predetermined 
amount (through appropriate preloading of the accumulator). Successive 
samples occurring after the next successive occurrence of the other of the 
high down or low up pulses are compared to the stored count and the 
trigger signal generated when the sampled Z counts favorably compare with 
the stored count. It should be appreciated that the amplitude, slew rate, 
Z out-of-limits and time domain discrimination are inherent in the X-ray 
triggering since the X-ray triggering derives from the high down and low 
up pulses. In addition, it is requisite that the X-ray machine be armed 
through an operator button, and the respiration rate must be constant 
within predetermined limits, e.g., varying less than .+-.20 breaths per 
minute. 
Referring now to FIG. 7, there is shown a suitable maxima X-ray trigger 
circuit 700. In practice, multiplexing techniques are utilized such that 
the same circuitry performs both maxima triggering and minima triggering. 
However, for the sake of clarity, only maxima triggering will be addressed 
at this point. 
Triggering circuit 700 is optically coupled to the X-ray machine through 
line receiver 702. An arm cycle is initiated when the operator depresses a 
spring loaded switch associated with the expose control of the X-ray 
machine. If, at any time prior to actual triggering, the switch is 
released the arm cycle will be inhibited. Thus, primary control of the 
X-ray machine is left in the hands of the operator. 
When the operator's arms switch is depressed, line receiver 702 generates a 
high level signal to arm logic 704. Arm logic 704 suitably comprises, in 
effect, AND gate logic to ensure that, (1) the arm button is depressed, 
(2) that no alarm condition exists (as will be explained), and (3) that 
the respiratory waveform period variations are within predetermined limits 
(e.g., .+-. a predetermined percentage or equivalent count). The period 
variation determination is provided by respiratory tachometry circuit 600. 
Arm logic 704 also includes a synthetic arm provision for testing the 
triggering apparatus in the absence of a connection to an X-ray machine. 
Arm logic 704, when enabled, generates an arm signal at terminal 704a. The 
arm signal at terminal 704a is applied to the clear (cl) terminals of 
respective JK flip-flops 714, 718, 728, 730, 746 and 748, each flip-flop 
having J input tied high and K input tied low. The arm signal at terminal 
704a is also applied as one input signal to a conventional 4-input AND 
gate 742 and is inverted and applied, through a 2-input OR gate 762, to a 
panel light rearm display 764. Thus, when the arm button is initially 
depressed and the system is in armed status, per arm logi 704, flip-flops 
714, 718, 728, 730, 746 and 748 are initialized to zero. 
Flip-flop 714 is clocked by the high down pulses generated at terminal 556a 
by one-shot 556. The Q output of flip-flop 714 is coupled to one input of 
a 3-input AND gate 716. The other input terminals of AND gate 716 are 
responsive to the ZCEP3 (reset) signal generated at terminal 408c of 
timing logic 408, and the Q output of flip-flop 718. 
The output of AND gate 716 is coupled to the reset terminal of a 16-bit 
binary counter 732, to the clock terminal of `D` flip-flop 730 and, 
through a delay 722 to the clock terminal of `D` flip-flop 718. Counter 
732 is clocked by the output of a 3-input NAND gate 720. 
The inputs of NAND gate 720 are receptive of the Z counts from terminal 
436a of digital peak detector 404, the Q output of flip-flop 730 and the Q 
output of flip-flop 728, respectively. The parallel outputs of counter 732 
are coupled to one set of input terminals of a magnitude comparator 734. 
The second set of input terminals at magnitude comparator 734 are coupled 
to the parallel outputs of a second 16-bit binary counter 738. 
Binary counter 738 is clocked by the output of a Q input NAND gate 736. 
NAND gate 736 is responsive to the Z count signal from terminal 436a of 
peak detector 404 and to the Q output of flip-flop 728. The clock input of 
flip-flop 728 is responsive to the output of a 2-input AND gate 726. AND 
gate 26, in turn, is responsive to the ZCEP2 signal generated at terminal 
408b of timing logic 408 and to the Q output of flip-flop 718. 
As previously noted, the arm signal from arm logic 704 is applied as an 
input signal to a 4-input AND gate 742. A second input of AND gate 742 is 
responsive to the outputs of a 2-input AND gate 740. AND gate 740 is 
responsive to the A less than B (A&lt;B) output signal of magnitude 
comparator 734 and to the ZCEP1 control signal generated at terminal 408a 
of timing logic 408. The high down to low up 222 msec. pulse generated by 
one-shot 562 is responsive to generation of the high down pulse and is 
inverted by an inverter 744 and applied to a third input of AND gate 742. 
The fourth input of AND gate 742 is responsive to the Q output of 
flip-flop 746. 
The clock terminal of flip-flop 746 is coupled to the output of a 2-input 
NAND gate 748. The inputs of NAND 748 are the low up pulses generated by 
one-shot 554 at terminal 554a and the Q output of flip-flop 714. The 
output of NAND gate 742 is applied to one input terminal of a further 
3-input NAND gate 756. The other inputs of NAND gate 756 are responsive to 
the inverted (through inverter 754) output signal of a shift register 752 
and to the Q output of JK flip-flop 748. Flip-flop 748 is clocked by the 
output of AND gate 756. The Q output of flip-flop 748 is coupled to a 
display lamp "exposed" on the indicator panel. Flip-flop 748 is initially 
cleared by the arm signal. Thus, when a trigger signal is generated, the Q 
output goes high to activate the "exposed" indicator and the Q output goes 
low, thereby inhibiting generation of additional trigger pulses. Shift 
register 752 is responsive to the high down pulse and is generated by 
one-shot 556 at terminal 556a. The output of shift register 752 is also 
applied as the second input to OR gate 762 to drive the rearm indicator 
764. The output of NAND gate 756 is applied to a line driver 450 to 
optically couple a trigger signal to the X-ray unit. 
In operation, assuming the system to be in armed status, when the operator 
depresses the arm button, the next successive high down pulse clocks 
flip-flop 714. The Q output assumes a high state upon the falling edge of 
the pulse. Recalling that flip-flop 718 was initially cleared, NAND gate 
716 thus generates a high level pulse upon the next occurrence of control 
signal ZCEP3 (reset) from timing logic 408. The pulse generated by AND 
gate 716 resets counter 732. 
Counter 732 is preloaded with a count corresponding to 0.3 ohms. The 
preloading of the count equivalent to 0.3 ohms, in effect, adds the count 
equivalent to 0.3 ohms to the Z count at the high down instant, to account 
for the resolution of the system. As previously explained, the high down 
(or low up) pulses are generated a maximum of 0.3 ohms after the 
occurrence of the peak. Thus, by adding the preload count to counter 732, 
the stored value corresponds to a point on the respiration waveform just 
preceding the inspiration maxima. 
The ZCEP3 pulse gated by AND gate 716 also clocks flip-flop 730, and upon 
the negative going transition of the pulse a high level Q output signal is 
generated to NAND gate 720. Recalling that flip-flop 728 was initially 
cleared, counter 732 thus accumulates the Z count associated with the 
sample immediately after the high down point. It should be appreciated 
that in view of storing the sample count occurring just subsequent to the 
high down point, the A less than B signal will be generated slightly 
before the maxima. Thus, time is provided for the X-ray machine to fire 
after generation of the trigger signal. 
After the delay by delay 722, flip-flop 728 is also clocked and AND gate 
726 enabled with respect to the next successive ZCEP2 (load) signal 
generated by timing logic 408 (terminal 408b). Flip-flop 728 is thus set, 
inhibiting NAND gate 720 and enabling NAND gate 736. Accordingly, counter 
738 accumulates the Z count of the next sample. The count in counter 738 
is compared to the count in counter 732 by comparator 734. Counter 738 is 
thereafter reset by the ZCEP3 pulse generated at terminal 408c of timing 
logic 408 in preparation for accumulation of the next sample Z count. 
In effect, each successive sample is compared to the count stored in 
accumulator 732 until the respiratory waveform reaches the point 
corresponding to the stored count in the next respiratory cycle. When 
counter 738 accumulates a count greater than the count stored in counter 
732 a signal is generated to enable AND gate 740 with respect to the ZCEP1 
timing signal (generated at terminal 408a of timing logic 408). Assuming 
that the output pulse from AND gate 740 is not generated during the high 
down to low up period (as determined by one-shot 562a), and that a low up 
pulse has occurred during the interum (as determined by NAND gate 748 and 
flip-flop 746), a signal is generated to NAND gate 756. NAND gate 756 is 
enabled to trigger the X-ray pulse through line driver 450 if the X-ray 
has not previously been exposed (as determined by JK flip-flop 748), and 
no more than a predetermined number (e.g., 5 breaths) have passed since 
the initial arming of the system (as determined by shift register 752). 
However, once shift register 752 has indicated that 5 successive breaths 
have passed during the arm sequence, or flip-flop 748 has indicated that 
the X-ray film is exposed (i.e., the picture has already been taken), AND 
gate 756 is inhibited until such time as the system is rearmed. Thus, only 
a single X-ray picture will be taken during the course of an arm sequence 
irrespective of the arm button being depressed for a length of time after 
the X-ray has been taken. Further, an indication of a malfunction is 
provided if 5 breaths pass without a triggering, and rearming of the 
system is required. 
Minima triggering would be established in a similar manner. However, the 
low up pulse generated at terminal 554a would be applied as a clock to 
flip-flop 714 and to shift register 754. Similarly, the high down signal 
generated at terminal 556a would be applied to AND gate 748, and the low 
up to high down one-shot pulse generated at terminal 558would be applied 
through inverter 744 to AND gate 742. In addition, counter 738, rather 
than counter 732, would be preset with the count equivalent to 0.3 ohms. 
In practice, the 0.3 ohm preload is hardwired into counter 732 and a count 
equivalent ot 0.6 ohms selectively preloaded into counter 738 through the 
multiplexers. The effect of the 0.6 ohm preload is to offset the hardwired 
0.3 ohm preload in counter 732 and subtract and additional 0.3 ohms from 
the count stored in counter 732 (effected during the comparison). 
As previously noted, preamplifier 200 generates an analog signal indicative 
of the biopotential of the myocardia, commonly referred to as an ECG 
signal. The ECG signal is differentially provided at terminals 238a and 
238b. A rate analysis similar to that performed on respiration is made on 
the ECG signal. 
The ECG signal is applied to an ECG analog processing circuit 800 
comprising a conventional QRS detector one-shot and time domain 
discrimination circuitry. ECG analog processing circuit 800 is shown in 
FIG. 8. The ECG signal provided at preamplifier output terminals 238a and 
238b is applied to a conventional QRS detector one-shot 802. QRS detector 
802 provides a 240 msec. pulse in response to each QRS complex. The QRS 
detector pulse train is applied to a pulse width modulator 804. Pulse 
width modulator 804 generates an output pulse at terminal 804a having a 
pulse width inversely related to the instantaneous heart rate. Pulse width 
modulator 804 comprises a 555 timer integrated circuit responsive to an 
analog signal indicative of the heart rate derived, as will be explained, 
from cardio-tachometry circuit 900. 
The pulse width modulated signal is applied to cardio-tachometry circuit 
900 shown in block diagram form in FIG. 9. Cardio-tachometry circuit 900 
is substantially identical to the respiration tachometry circuit 600 
except that the timing chain 908 (analogous to timing chain 608) generates 
respective output signals at twice the rate of the analogous signals in 
respiration tachometry circuit 600: 960 pps (used instead of 480 pps), 60 
pps (used instead of 30 pps), and 30 pps (used instead of 15 pps). 
Further, the convert period generated by a shift register 910 analogous to 
shift register 610 is 33 msec. in length, rather than 66 msec. as in the 
respiration tachometry circuit 600. The preload count in the binary 
counter in a period to rate converter 920, corresponding to counter 614 in 
period to rate converter 620 of respiration tachometry circuit 600 
therefore differs accordingly. In addition, the averaging function 
provided by selective averaging logic 924 analogous to selective averaging 
logic 624 provides for averaging over 10 heart beats rather than 5 breaths 
as in respiration tachometry circuit 600. Accordingly, the counters 
corresponding to counters 626 and 630 are divide-by-10 counters rather 
than divide-by-5 counters. In addition, in the circuit 964 corresponding 
to artifact rejection circuit 664 components corresponding to AND gate 694 
would be deleted such that the components corresponding to OR gate 671 is 
responsive to two input signals, A less than B and force (from the 
magnitude comparator and shift register). Thus, heart rate counts 
differing by .+-.20 beats per minute from the base rate are rejected. The 
output of the D/A converter responsive to the outputs of the latches 
corresponding to latches 662 in artifact rejection circuit 664 is utilized 
as the analog heart rate signal to time domain discrimination pulse width 
modulator 804. 
It will be understood that when an element of the cardio-tachometry circuit 
900 is hereinafter referred to, the designation will be the last two 
digits of the corresponding element of respiration tachometry circuit 600 
prefaced with a 9 rather than a 6. For example, the counter in 
cardio-tachometry circuit 900 corresponding to counter 660 of respiration 
tachometry circuit 600 will hereinafter be referred to as counter 960 and 
so forth. Similarly, analogous control signals in cardio-tachometry 
circuit 900 will be designated by the prefix H, replacing the R used with 
respect to tachometry circuit 600. For example, the signal in 
cardio-tachometry circuit 900 analogous to the signal RCEP1 will be 
designated HCEP1, and so forth. 
As previously noted, monitor 100 generates an apnea alarm in response to a 
number of different detected conditions. An alarm is generated if the 
respiration period extends beyond a predetermined length, as detected by 
the generation of the signal R overflow by period counter 614 at terminal 
614a. It should be appreciated that other bits of counter 614 besides the 
most significant bit, can be used to initiate an alarm. In the preferred 
embodiment, an alarm is generated if the respiration period exceeds 16 
seconds. A similar alarm is provided in cardio-tachometry circuit 900 with 
respect to heart beat rate. The alarm is generated if the heart beat 
period exceeds 1.8 seconds. 
An apnea alarm is generated if a physiological apnea occurs, that is, if 
the respiration period doubles in successive samples and four successive 
heart beat decelerations are detected. Physiological apnea detector 1000 
will be hereinafter more fully described in conjunction with FIG. 10. 
Referring now to FIG. 10a, circuit 1000 utilizes two identical banks of 
calculation apparatus, bank A and bank B. Calculation apparatus bank A 
comprises a preloaded (16 counts) binary counter 1002a, a preloaded (8) 
binary counter 1004a, a divide-by-2 counter 1006a, a magnitude comparator 
1008a, and two 2-input AND gates 1010a and 1020a. Divide-by-2 counter 
1006a is coupled to the clock input of binary counter 1004a. One input of 
2-input AND gate 1010a is coupled to the A&gt;B output of magnitude 
comparator 1-08a, and one input of 2-input AND gate 1020a is coupled to 
the A&lt;B output of comparator 1008a. 
Calculation apparatus bank B is essentially identical to calcuation 
apparatus bank B and comprises a preloaded (16) binary counter 1002b, a 
preloaded (8) counter 1004b, a divide-by-2 counter 1006b, a magnitude 
comprator 1008b, and 2-input gates 1010b and 1020b. 
In operation, suitable timing logic 1014, illustrated diagrammatically in 
FIG. 10b, generates three sequential timing signals, EP1a on terminal 
1014a, EP2 on terminal 1014b and EP3 on output terminal 1014c. 
Suitable steering logic 1012 (shown diagrammatically in FIG. 6c) 
alternately applies control signals generated by timing logic 1014 to the 
respective calculation apparatus banks A and B such that the period count 
is accumulated in the respective counter 1002 of a first bank during a 
first cycle, and such that during the next cycle, the period counts are 
applied through the respective divider 1006 to the respective binary 
counter 1004 of the first bank, while the period count is simultaneously 
accumulated in the binary counter 1002 of the other bank. 
For example, assuming bank A to be initially addressed, steering logic 1012 
provides the timing signals at the A output terminals thereof. A first 
period count would be accumulated in binary counter 1002a. The next period 
count generated would be applied simultaneously through divider 1006a 
binary counter 1004a and to binary counter 1002b. Thus, at the end of the 
second count, binary counter 1002a contains a first period count and 
binary counter 1004a contains 1/2 of the second period count and counter 
1002b contains the second period count. 
The contents of binary counters 1002a and 1004a are compared by magnitude 
comparator 1008a. If the contents of counter 1004a (1/2 the second period 
count) are greater than the previous count contained in counter 1002a, the 
second period is twice the duration of the preceding period. Accordingly, 
the A greater than B signal enables AND gate 1010a with respect to a 
EP1 timing pulse generated by timing logic 1014 provided at steering 
logic terminal DMUX1a. The EP1 pulse is gated through an OR gate 1016 
to clock a D flip-flop 1018. The D input of flip-flop 1018 is tied high, 
and accordingly, a high level Q output is provided. If A is less than B, 
that is, the second period count is not twice the previous period count, 
AND gate 1020a is enabled with respect to the EP1 pulse (provided at 
steering logic terminal DMUX1a) to clear flip-flop 1018. 
In the next cycle, steering logic 1012, in effect toggles, providing the 
control signals on the B output terminals thereof to reverse the roles of 
calculation apparatus bank A and calculation apparatus bank B. 
Accordingly, the period count is applied through divide-by-2 counter 1006b 
to binary counter 1004b and is simultaneously accumulated in binary 
counter 1002a. Thus, at the end of the conversion, counter 1002b contains 
the previous period count, counter 1004b contains 1/2 of the instantaneous 
period count, and counter 1002a contains the instantaneous period count. 
The contents of counters 1002b and 1004b are compared by comparator 1008b 
and the A greater than B or A less than B signals generated accordingly. 
As previously noted, when an A&gt;B signal is generated, flip-flop 1018 
assumes a high level Q output. The Q output is applied to a 1 .mu.sec. 
one-shot 1026, initiating circuitry for detection of four successive 
decelerations in heart rate. Recalling that the up/down control flip-flop 
976 for the up/down counter 972 in the artifact rejection portion 964 of 
cardio-tachometer 900 provides an indication of whether the heart rate has 
increased or decreased with respect to the previous period, a high 
Hcountup (Q) signal, supplied at terminal 976a, is indicative of an 
increased heart rate. Similarly, an Hcountdown signal (Q), supplied at 
terminal 976b is indicative of a decreased heart rate. Accordingly, the 
Hcountup signal (976a) is gated in an AND gate 1028 with the HCEP1a-1 
signal (indicative of a valid rate count) provided at terminal 966a in the 
cardio-tachometer 900 and the gated signal applied along with the output 
of one-shot 1026 to the respective inputs of a 2-input NOR gate 1030. The 
output of NOR gate 1030 is coupled to the reset terminal of a shift 
register 1032. NOR gate 1030 generates a low level signal in response to a 
high level signal from either one-shot 1026 or AND gate 1028. 
In operation, the one-shot pulse triggered by the Q output of flip-flop 
1018 initially resets shift register 1032. Shift register 1032 is clocked 
by the output of a 2-input AND gate 1034 which gates the Hcountdown signal 
from terminal 976b of cardio-tachometer 900 with the HCEP1a-1 (valid rate 
count) signal from terminal 966a. Thus, shift register 1032 is indexed by 
each decreasing heart rate sample. However, shift register 1032 is reset 
in response to a high level output from AND gate 1028 generated in 
response to an increasing heart rate (Hcountup). Thus, if four decreasing 
heart rate samples occur without an intervening increasing heart rate 
sample shift register 1032 generates at its QD output a high level signal 
which is applied, along with the Q output of flip-flop 1018 to a 2-input 
AND gate 1036. The output of NAND gate 1036 is applied as a clock input to 
a D flip-flop 1038. The D input of flip-flop 1038 is tied high, and the Q 
output is tied to one input of a 2-input OR gate 1040. The other input of 
OR gate 1040 is responsive to the R overflow signal from period counter 
terminal 614a (respiration tachometry circuit 600). Accordingly, OR gate 
1040 generates a high level signal if a respiration period is twice the 
duration of the preceding period and four valid decreasing heart beat 
samples are detected or if the period counter 614 overflows indicating a 
period beyond the maximum threshold. The high output of OR gate 1014 
initiates an apnea alarm. 
As noted above, an apnea alarm is also generated as a fail-safe, in the 
event that the detected heart rate and respiration rate are approximately 
equal for a predetermined number of successive samples. This condition is 
detected by cardiac component detection sensor circuit 1100 and will now 
be described more fully in conjunction with FIG. 11. In general, sensor 
1100 captures an RPDO count indicative of the respiration rate in breaths 
per minute and compares it to an HPDO, the analogous numerical number 
equivalent to the heart rate in beats per minute. However, the RPDO counts 
and HPDO counts are asynchronous data. Accordingly, an RPDO count is 
captured at random and held, then counted down using timing derived from 
cardio-tach 900. The captured RPDP count is counted down with the next 
successively occurring HPDO and if more than four successive differences 
are within a predetermined limit, an alarm is generated. 
Referring now to FIG. 11a, the RPDO signals from period to rate converter 
620 (terminal 619a) are applied to one input of a 3-input NAND gate 1116, 
the output of which is applied as a clock input to an 8-bit binary counter 
1102. The parallel outputs of counter 1102 are applied to the inputs of an 
up/down counter 1124, the parallel outputs of which are in turn coupled to 
a first set of inputs of a magnitude comparator 1132. The second set of 
inputs to comparator 1132 have applied a hardwired preloaded count 
corresponding to the desired limit. The max/min output of up/down counter 
1124 (indicative of the counter attaining a zero count) is applied to 
clock an up/down control flip-flop 1138 in a manner similar to the 
cooperation between up/down counter 672 and flip-flop 676 in artifact 
rejection circuit 664 (respiration tachometry circuit 600). 
Another input of NAND gate 1116 is responsive to the Q output of a JK 
flip-flop 1118. The J and K inputs of flip-flop 1118 are respectively tied 
high and to ground. JK flip-flop 1118 is clocked by the output signal of a 
2-input AND gate 1112. AND gate 1112 is responsive to the Rload signal 
generated at timing logic output terminal 612d and to the Q output of D 
flip-flop 1114. The third input of NAND gate 1116 is also responsive to 
the Q output of flip-flop 1114. The D input of flip-flop 1114 is tied 
high. Flip-flop 1114 is clocked and binary counter 1102 reset in response 
to a signal generated by a 3-input AND gate 1106. 
AND gate 1106 is responsive to the RCEP signals generated at timing logic 
output terminal 612c, to the Q input of a JK flip-flop 1104 and to the Q 
output of a D flip-flop 1108. JK flip-flop 1104 and J and K input tied 
high and to ground, respectively, and is clocked by the RECEP1a-1 signal 
generated by artifact rejection circuit 664 at terminal 666a. It should be 
recalled that the RCEP1a-1 signal is indicative of the accumulation of a 
valid rate count. D flip-flop 1108 and D input tied high and is clocked by 
the output signal of AND gate 1106, delayed by a predetermined delay 1110. 
The Q output of flip-flop 1118, the Q input of flip-flop 1114 and the 
Rload signal generated at timing logic output terminal 316d are applied to 
the respective inputs of a 3-input NAND gate 1122. The output of NAND gate 
1122 is applied to the reset terminal of up/down counter 1124. 
The Q output of flip-flop 1118 is also applied to one input of a 2-input 
AND gate 1126. The other input of AND ate 1126 is the HCEP3 signal 
generated at the cardio-tachometry timing logic output terminal 912c. The 
output of AND gate 1106 is applied to the clock input of a D flip-flop 
1128. 
The D input of flip-flop 1128 is tied high. The Q output of D flip-flop 
1128 is applied to one input of a 2-input NOR gate 1130 and to one input 
of each of 3-input AND gates 1134 and 1146. The other input of NAND gate 
1130 is the HPDO counts from the cardiotachometry circuit period to rate 
converter input terminal 919a. 
Second input terminals of each of AND gates 1134 and 1136 are responsive to 
the RCEP1a signal generated at respiration tachometry circuit timing logic 
output terminal 612e. The third terminals of AND gates 1134 and 1136 are 
respectively responsive to the A&lt;B and A&gt;B output signals of magnitude 
comparator 1132. The output of AND gate 1134 is applied as the clock 
signal to a shift register 1138. The QC output (corresponding to the third 
bit) of shift register 1138 provides an activation signal for respective 
audio/visual alarms) and is used to set the respiration display to zero. 
The output of AND gate 1136 is applied to one input of a 2-input NOR gate 
1142. The other input to NOR gate 1142 is the master clear signal 
generated at terminal 316a of analog processing circuit 300 (FIG. 3a). The 
output of NOR gate 1142 is applied as a clear signal to shift register 
1138 and to one input of a 2-input OR gate 1137. The other input of OR 
gate 1137 is responsive to the output of AND gate 1134. The output of NOR 
gate 1137 is applied to one input of a 2-input NOR gate 1139, the other 
input of which is responsive to the master clear signal from terminal 
316a. The output of NOR gate 1134 (hereinafter designated reset) is 
applied as a reset signal to flip-flops 1104, 1108, 1114, 1118 and 1128. 
In operation, the circuit initialized in response to the master clear 
signal, and the respective flip-flops of circuit 1100 reset. When 
RCEP1a-1, indicative of the accumulation of a valid respiratory count is 
generated, flip-flop 1104 generates a high level Q signal. Accordingly, 
since D flip-flop 1108 is initially reset, AND gate 1106 is enabled with 
respect to the next occurring RCEP3 pulse. Accordingly, binary counter 
1102 is cleared and flip-flop 1114 set. 
The high level Q output of flip-flop 1114 enables NAND gates 1116 and 1122, 
(since flip-flop 1118 is initially reset) with respect to the RPDO signal 
and Rload signal, respectively. The high Q output from flip-flop 1114 also 
enables AND gate 1112 with respect to the next occurring Rload signal. 
Accordingly, the next occurring Rload signal clocks JK flip-flop 1118, 
causing the Q output to assume a high level and the Q output a low level 
at the falling edge of the Rload pulse. NAND gates 1116 and 1122 are thus 
inhibited at the falling edge of the Rload pulse. The RPDO counts are, 
however, accumulated in binary counter 1102 during the interim between the 
RCEP3 pulse and the occurrence of the Rload pulse. The Rload pulse, in 
effect, causes up/down counter 1124 to be loaded with the contents of 
binary counter 1102. 
On the falling edge of Rload, AND gate 1126 is enabled with respect to the 
HCEP3 ;signal. Accordingly, the next successive HCEP3 signal clocks D 
flip-flop 1128 to initiate a down count operation. The high Q output of 
flip-flop 1128 enables AND gate 1130, causing the HPDO counts (indicative 
of heart rate) to count down up/down counter 1124. At the end of the count 
down sequence, the contents of up/down counter 1124 represent the 
difference between the heart rate and respiration rate. 
The difference count is then compared to the preloaded threshold (e.g., 4) 
by magnitude comparator 1132. If the difference is greater than the 
preloaded count, the A&gt;B signal enables AND gate 1136 with respect to the 
RCEP1a signal, causing NOR gate 1142 to generate a low level signal to 
reset shift register 1138. Conversely, if the difference is less than the 
preloaded count, the A&lt;B signal from comparator 1132 enables AND gate 1134 
with respect to the RCEP1a signal causing shift register 1138 to be 
indexed. The passage of the RCEP1a pulse by either AND gates 1134 or 1136 
effects a reinitialization of the system in readiness for the next 
respiration rate count. 
If three successive comparisons results in A&lt;B signals, without an 
intervening A B comparison, shift register 1138 actuates an alarm. The 
respective signals associated with circuit 1100 are illustrated in timed 
relation in FIGS. 11b and 11c. 
As previously noted, it is desirable to determine the number of apneic 
episodes occurring during a given time period. A circuit 1200 for 
providing an indication of apneic episodes per unit time is shown in FIG. 
12. A clock 1202 operating from and in synchronism with the power line 
signal generates a one pulse per second signal. The one pulse per second 
signal is applied to a 2-input AND gate 1204. 
AND gate 1204 gates the one pulse per second signal with the output (valid 
time) of a 2-input AND gate 1206. AND gate 1206 is responsive to the 
master clear signal from terminal 316a and the arithmetic enable signal 
from terminal 410a. The output of AND gate 1206 provides an indication of 
the actual time during which a valid monitoring of the subject is provided 
by monitor 100. The output of AND gate 1204 is applied to a divide-by-60 
BCD counter chain 1208. The most significant bit of divide-by-60 counter 
chain 1208 is applied to a 500 msec. one-shot 1210. 
The one-shot pulse is applied to suitable timing logic 1212. Timing logic 
1212 generates a response to the one-shot pulse three sequential output 
signals, in sequence: load count, latch count and reset. 
The valid time signal is also applied to one input of a 2-input AND gate 
1214. The other input of AND gate 1214 is connected to a rotary switch 
1216 which selectively couples the input to one of three terminals, 1216a, 
1216b and 1216c. Terminal 1216a provides an indication of an apnea 
generated by a 3-input OR gate 1218. OR gate 1218 is responsive to the 
physical apnea signal from OR gate 1040, the R overflow signal provided at 
terminal 614a and the apnea/low rate indication provided by the Q output 
of flip-flop 647. Terminal 1216b is connected to terminal 967a to provide 
an indication of bradycardia. Terminal 1216c provides an indication of any 
apnea and bradycardia by coupling the respective terminals 1216a and 1216b 
to the respective inputs of a AND gate 1220. Thus, AND gate 1214 provides 
a pulse in response to any occurrence during valid monitoring time of the 
event chosen by switch 1216. 
The output of AND gate 1214 is applied as a clock input to a BCD counter 
1222. The parallel outputs of BCD counter 1222 is coupled to a register of 
latches 1224, the parallel outputs of which are in turn coupled to a 
parallel-in-serial-out shift register 1226. The output of shift register 
1226 is applied to a seven segment display 1228. BCD counters 1222 
accumulate a count representative of the number of occurrences of the 
chosen event. 
In operation, in response to the 500 msec. one-shot pulse, timing logic 
1212 generates control signals to cause shift register 1226 to be loaded 
with the contents of latches 1224. Latches 1224 are then updated with the 
contents of BCD counters 1222. Counters 1222 are then reset for 
accumulation of the number of events occurring during the next period of 
time. 
It should be understood that the above description is of illustrative 
embodiments of the present invention, and that the invention is not 
limited to the specific form shown. Various modifications can be made in 
the design and arrangement of the elements as will be apparent to those 
skilled in the art without departing from the scope of the invention as 
expressed in the appended claims.