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Publication numberUS3572317 A
Publication typeGrant
Publication dateMar 23, 1971
Filing dateOct 2, 1968
Priority dateOct 2, 1968
Also published asDE1949445A1
Publication numberUS 3572317 A, US 3572317A, US-A-3572317, US3572317 A, US3572317A
InventorsGerald James Wade
Original AssigneeHoffmann La Roche
Export CitationBiBTeX, EndNote, RefMan
External Links: USPTO, USPTO Assignment, Espacenet
Respiratory distress monitor
US 3572317 A
Abstract  available in
Previous page
Next page
Claims  available in
Description  (OCR text may contain errors)

March 23, 1971 G, J, WADE RESPIRATORY DISTRESS MONITOR 4 Shee's-Sheet l Filed OCT.


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HBLBIN LV 8 HOLVHOELLNI :40 lndlHO B LnNlW/SLVBE-.LHVBH United States Patent O 3,572,317 RESPIRATORY DISTRESS MONITOR Gerald James Wade, Clifton, NJ., assgnor to Hoffmann- La Roche Inc., Nutley, NJ. 'Filed Oct. 2, 1968, Ser. No. 764,448 Int. Cl. A61b 5 02, 5 08 U.S. Cl, 12S-2.05 17 Claims ABSTRACT OF THE DISCLOSURE The invention relates to an apparatus for accurate determination of respiratory distress conditions, especially those often experienced by newborns.

One of the most serious problems in pediatrics today is the large number of neonate fatalities and mentally retarded infants resulting directly or indirectly from the phenomena of respiratory distress syndrome, which in brief might be defined as the inability of the newborn infant, regardless of cause to make the necessary respiratory adaptation following birth. The exact cause(s) of this phenomena are not altogether known and could only be ill-defined in part because the extreme variations in physiologic conditions of the well adapted newborn, however, the belief is advanced by many that one of the most important causes of respiratory distress in the newborn involves the lung tissue itself and the formation of a hyaline membrane in this tissue. When this latter condition is present, breathing can become erratic and often stops for intervals of time commonly referred to as respiratory distress. The result is undesirable changes in blood gases and corresponding physiologic damage to other tissues.

As in other medical areas, advances have been made in surgical and medicinal means for alleviating and eliminating respiratory distress conditions, and in fact, the most well known and commonly used curative technique consists of applying a slight mechanical stimulus to the neonate by the nurse, which usually suffices to restore normal breathing. However, in order to timely apply suitable curative techniques it is extremely important to immediately detect the onset of a respiratory distress attack to avoid or minimize the occurrence of physiologic damage or prevent a fatal mishap. As a practical matter, however, it is unrealistic to expect virtually constant observation of newborns even in a well organized intensive care unit. Although extreme vigilance, which is required in nursing staffs in such units, may generally be sufficient to detect a respiratory distress condition in time to prevent the occurrence of a fatality, it may not be early enough to prevent physiological failings which may eventually cause retardation or other physiologic damage. Accordingly, a monitoring system adapted to continuously observe a neonate and provide an early warning alarm at the onset of a respiratory distress condition to enable early corrective action to be taken, would be highly desirable.

Monitoring devices of the type referred to above are commercially available, whereby basically these devices attempt to provide signals for respiratory distress conditions by indiscriminately analyzing respiratory activity of the neonate. Unfortunately, however, newborn infants breath very irregular, whereby non-critical respiratory arrests are a common occurrence in such irregular breathing. Similar respiratory arrests heretofore referred to as respiratory distress may arise which might lead to total respiratory failure and death or a prolonged respiratory failure and subsequent physiologic damage such as brain damage or physical handicap, if not detected early by the medical staff. In View of the above, and for the fact that respiratory rates and short respiratory arrests are poor indicators of the severity of respiratory distress, commercial neonatal monitors basically utilizing respiratory activity alone for detecting respiratory distress co'nditions have been found ineffective for the reason that they do not clearly distinguish incipient respiratory distress conditions from non-critical respiratory arrest conditions, and thus too often give false indications of respiratory distress causing severe loss of nursing time and frequent inaccurate records leading to needless worry and precautions.

It is the purpose of the present invention to overcome the shortcomings of the prior art by providing a monitor adapted to distinguish between non-critical respiratory arrest conditions and critical respiratory arrest conditions. The latter has been accomplished by the discovery of a heart beat activity and respiratory activity interaction and specifically that pronounced bradycardia and respiratory arrests when taken together and analyzed, sharply delineate the newborn requiring immediate attention from the child who is varying his routine a little with temporary non-critical respiratory arrests and is not in any immediate danger. In implementing the above, prescribed minimum heart rate values and prescribed minimum respiratory arrest time values are related to one another relative to time. Aside from the above discovery it has additionally been found that in certain instances of abnormally extended respiratory arrest resulting in a respiratory distress condition, bradycardia is absent, and by providing a detection system for responding only to extremely low respiration rates to supplement the detecting of the above discussed interactions, most all conditions of respiratory distress can be continuously monitored with virtually little chance of causing false alarms.

Other objects, advantages and capabilities of the present invention will become apparent from the following detailed description, taken in conjunction with the accompanying drawings, showing only preferred embodiments of the invention.

I n the drawing:

FIG. l is a block diagram of one form of the present invention.

FIG. 2 represents a detailed schematic block diagram of the invention.

FIG. 3 illustrates a series of time related voltage waveforms corresponding to points indicated in FIG. 2.

FIG. 4 pictures a recorded comparison of the beat-tobeat heart activity signal and an output signal display represented by heart-beats per minute of the integrated heart-beat value over eight second periods.

FIG. 5 displays representations of respiratory rate and Iheart rate values as would resceptively appear on meters 47 and 28 in FIG. 2.

GENERAL DESCRIPTION Turning iirst to FIG. 1 there is shown a block diagram of a general monitoring system utilizing the principles of this invention, wherein a patient 11 is connected by suitable transducers (not shown) to blocks 12, and 13 for provding representative electrical signals to respective- 3 ly measure heart-beat activity and respiratory activity parameters of the patient, which parameters are used as a base lfrom which a condition of respiratory distress is evaluated, which respiratory distress will be particularly described with reference to neo-natals in the instant disclosure.

It has been found that by relating the physiological parameters of heart-beat activity and respiratory activity to one another as is performed in block 14 a condition of respiratory distress may be detected and distinguished from temporary breathing stoppages which are non-critical. This is quite a significant advancement in neo-natal monitoring for the reason, as explained above, because newborn premature infants do not breath consistenly but instead exhibit widely fluctuating rates and periods of irregular breathing including extensive breathing stoppages with no apparent ill effects.

In approaching a suitable method for relating the latter mentioned physiological parameters, it is rst observed that each of the specified physiological parameters to heart-rate data and to use the waveform of respiratory activity to provide respiratory arrest data as evidenced by blocks 15 and 16 respectively. Detection of prescribed low heart-rate events and minimum prescribed periods of cessation of breathing, applicable for neo-natals, are also employed in blocks 15 and 16 respectively, whereby simultaneous occurrence of such events are related in block 17 which would respond to denote a condition of respiratory distress and in turn activate an alarm 18 indicating that an infant requires immediate attention.

It is noted that the above as well as the ensuing description of the present invention has been tailored for use with neo-natal patients, however, it should be understood that the type of basic characteristics, i.e., waveform and rhythm, of the two physiological parameters employed as well as the prescribed circuit parameters i.e. frequency level sensor etc., could be altered to provide similar monitoring of respiratory distress victims not only with respect to neo-natals, but additionally to enable the invention to be equally suitable for use with older children and adult patients having symptoms involving respiratory distress conditions.

Further, in the ensuing description and referenced drawings it is to be noted that in addition to a detailed description for implementing detection of the degree of /respiratory distress above noted by relation of heart* beat to respiratory activity, apparatus is additionally disclosed to implement the detection of a second condition of respiratory distress indicating the need for immediate attention to the neo-natal patient, whereby the second condition of respiratory distress may be independently detected but serves to supplement the overall detection of respiratory distress conditions.

DETAILED EMBODIMENT Turning to FIG. 2, there is shown a schematic diagram of a respiratory distress monitoring system including the principles of this invention generally discussed above. Physiological electrodes 21 or other suitable transducers are positioned in contact with a patient 22 either directly or by way of an intermediate medium such as the conventional electrode pastes. Electrodes 21 are positioned on the patient to obtain a usable electrical signal representative of heart beat having the least artifact, such as that depicted in FIG. 3(a), with the conventional PQRST complex illustrated by the two successive heart beat electrical signals 23 and 24.

The signal from electrodes 21 is fed to a conventional differential amplifier 25 of the type commonly referred to as an ECG amplifier, having a high common mode rejection at sixty cycles and having a high input im- 4 pedance to not only minimize the flow of current through the electrodes, but additionally to compensate for the possibility of a pair of unbalanced electrodes. The output of the amplier is then directed to a lilter 26 designed to have a center frequency of 30 cycles and a bandwidth of about 10i cycles toy remove the P and T waves and pass the remaining signal whichlargely is the QRS complex, illustrated in FIG. 3(b). Although 30 cycles has been selected in the present embodiment as the center frequency of filter 26, it has been found that other center frequencies, i.e., 20 or 40 cycles are usable for obtaining a meaningful magnitude.

Connected to the output of lter 26 is a pulse Shaper unit 27 such as the Schmitt trigger circuit which will respond to the leading edge of the input signal 3(1)) to provide output pulses of uniform duration and amplitude such. as impulses 23' and 24' shown in FIG. 3(0) representative of each QRS complex picked up by electrodes 21.

The impulses from pulse shaper 27 are applied to the input of an integrator 28 of a conventional type such as an integrating amplifier, in which a potential increment is accumulated on a storage element (i.e. a condenser) proportional in value to the number of applied impulses representative of heart-beat from pulse shaper 27, to provide an output representative of heart-rate, which heart-rate is determined over a period equivalent to the RC time constant of integrator '28. The integration time constant (T) selected for the present embodiment is eight seconds, which time duration permits good meter displays of heart-beat rates from approximately thirty to two hundred beats per minute with a slowing rate in the neigborhood of ten beats per second. Such a time constant value has also been found sufficiently adequate to provide for ready observation of abnormal steady heart rates while at the same time virtually enabling normal varying heart-rates to be ignored, as these latter type heart rates become unnoticeable by being averaged out at such an integrating period (T). Although eight seconds has been selected for the present embodiment other integration time values for use `with neo-natals may be successfully utilized such as any value between, for example, five to twenty seconds.

Of course, it is understood, that the heart-beat electrical signals 23, 24 in FIG. 3(a), and similarly the impulses 23', 24 in FIG. 3(c) are but two of a long series or train of electrical signals picked-up during patient monitoring over an extended period of time. Assuming that the heart-rate of the patient being monitored, immediately prior in time to heart-beat impulse 23', is x heart-beats per period (T); then, in FIG. 3(d) the waveform 219 commencing at point x would be representative of the heart-rate output at integrator 28. It is observed that the trend of the heart-rate waveform 29 is positively altered at point y as at this point the potential of impulse 24 is charging the capacitor of the RC integrating time constant.

Connected at the output of integrator 28 are two units, one a meter 31 and the other a low level sensor 32. The meter 31 might be a galvanometer of the dArsonval type which for convenience, is adjusted to provide for a heart-rate reading in heart-beats per minute rather than heart-beats per an eight second period (T).

Sensor unit 32 is an adjustable threshold circuit adapted to respond to occurrences when the heart-rate of the patient being monitored falls below a certain prescribed frequency level or threshold based on the Period (T) at which the heart-rate is being evaluated as denoted by the output of integrator 28. Assuming a prescribed low heart-rate level is z as indicated in FIG. 3(d) by a dashed line, it is observed that waveform '29 will dip below the prescribed threshold z at crossing point v and remains below until crossing point w. As depicted in FIG. 3( e) sensor 32 will respond to heart-rate values at the output of integrator 2S below z to provide an output pulse 35 during the time duration between points v and w. The output of sensor 32 is then connected to one input of an AND gate 33. Several different type units may be utilized for sensor 32 such as a biased zero crossing detector which basically acts as a differential voltage comparator. In operation of the sensor unit 32 the voltage output signal of integrator 28 is compared with the reference voltage z to produce, in the present embodiment, a digital one, denoted as 33 in FIG. 3(e), lwhen the reference voltage input z is greater than the voltage output signal from integrator 28. A value of z found to be suitable would be equivalent to a low level frequency of about ninety heart-beats per minute.

In referring back to the patient, for obtaining electrical signals of respiratory activity there is schematically illustrated an impedance pneumograph arrangement including a pair of physiological electrodes 34 connected to a high frequency current source 35 to provide a high frequency carrier amplitude modulated respiration signal of the type disclosed in FIG. 3(1). Of course, impedance pneumography is only one of many techniques which may be ut1- lized in connection with the present invention for providing an electrical signal indicative of respiratory activity. Such other techniques might include the tidal volume approach or a strain gauge taped directly on the abdomen.

The respiration electrical signal representation 1s d1- rected from electrodes 34 to an A.C. differential amplifier 36 the output of which is fed to a high pass filter 37 for blocking out 60- signals resulting from the noisy environment. The filtered amplitude modulated signal is then supplied to a demodulator 38 to provide an output demodulated signal such as that depicted in FIG. 3(g), which signal is fed into a zero-crossing detector 39. As may be observed from FIG. 3(h), pulses of uniform amplitude and polarity are generated from the zero crossing detector in response to crossing of the zero axis by the output signal of demodulator 38. The output of zero crossing detector 39 is connected to a pulse shaper 41 such as a D.C. monostable triggered device which switches to its unstable condition for a predetermined time upon occurrence of an input signal exceeding the pre-selected trigger level thereby standardizing the random width output pulses of zero crossing detector 39 to provide output impulses 42 such as illustrated in FIG. 3(1').

One leg of the output lead of pulse shaper 41 is connected to an adjustable time delay one-shot multivibrator 43 which basically will generate an output impulse when the duration between successive input impulses 42 will exceed a predetermined time interval denoting absence of the breathing function or a respiratory arrest for at least the pre-set predetermined time interval. To attain the above, with relation to FIG. 3( j), the multi- Vibrator would include an adjustable RC circuit, the capacitor charging up at a known rate during the absence of input impulses and discharging upon the presence of an input impulse such as 42 whereby should the charging capacitor exceed a pre-adjusted lvoltage level n as shown at n', a predetermined time interval z reltaed to such voltage would be arrived at or exceeded prior to occurrence of a successive input pulse 42, and the multivibrator would typically respond to provide an output impulse 44 as illustrated in FIG. 3(k), until return of the breathing activity represented by a successive input impulse 42 appearing at the input of multivibrator 43. Suitable time periods for which 43 is adjustable have been found to be fifteen, thirty or forty-five seconds.

The output of multivibrator 43 is connected as a second input to AND gate 33 so that when positive output signals depicted at FIGS. 3(e) and 3(k) are simultaneously present at the input of AND gate 33 from the output of sensor 32 and multivibrator 43 to respectively denote a low heart rate and absence of breathing exceeding a pre-selected time interval, then AND gate 33 will be enabled to pass a signal through OR gate 45 and energize the alarm 46 to indicate the existence of a condition of respiratory distress. The alarm might either be of the audible or light signaling type, or both.

With reference back to pulse Shaper 41 the second leg of its output is connected to an integrator 47 similar to integrators 28 except that integrator 47, in the present embodiment is provided with an RC integrating time constant of thirty seconds to enable a good meter display of respiration rates from as low as fifteen beats per minute to 250 beats per minute. In FIG. 3(1), assuming that the respiration-rate of the patient being monitored immediately prior in time to impulses 42, is x breaths per period (thirty seconds); then, the waveform 48 commencing at point x would be representative of the respiration-rate output at integrator 47. It is noted that the trend of the respiration-rate waveform 48 is positively altered at point y' as at this point the potential of the second illustrated impulse in FIG. 3(1') is charging the capacitor of the RC integrating time constant.

Connected at the output of integrator 47 are two units, one a meter 49 and the other a low limit sensor 51. The meter, similar to meter 31, might be of the dArsonval type which is suitably adjusted to provide a respirationrate reading in breaths per minute as opposed to breaths per thirty second period.

Sensor unit 51 comprises a conventional adjustable threshold circuit similar in operation to sensor unit 32 hencetofore discussed whereby the sensor will be triggered when the respiration rate of the patient being monitored, based on the integrated output of integrator 47, falls below a certain prescribed frequency level, a Isuitable frequency level havin-g been found to be eighteen breaths per minute. For example, assuming the prescribed frequency level to be equivalent to a voltage level z' illustrated in FIG. 3(1), when the respirationrate waveform 48 fall below level z such as at point w', the sensor unit 51 will be triggered to provide au output Signal i.e., a digital one on lead 52. Whereas when Waveform 48 lies above voltage level y an output signal will appear at lead S3. Again either a Schmitt trigger a zero crossing detector, or similar suitable type unit may be utilized for sensor 51.

Output leads 52 and 53 of sensor unit 51 are connected to a suitable timer unit 54 having in the preferred ernbodiment an RC time constant of sixty seconds, the timer cycle, initiated by charging of the capacitor, being enabled by an output signal from sensor 51 on lead 52, and the timer adapted for reset by an output signal from sen-sor 51 on output lead 53 to cause the capacitor to discharge. If timer 54 is not reset by the end of a sixty second period the timer will be enabled to provide an Output signal 55 as depicted in FIG. 3(m) which signal is passed through OR gate 45 to trigger alarm 46.

OPERATION In reviewing the operation of the present invention, upon receiving from the patient heart-beat activity signals such as those illustrated in FIG. 3(a) the signals are amplified and filtered by ECG amplifier 25 and lter 26 respectively, and then translated into pulses of equal amplitude and width as depicted in FIG. 3(c). Simultaneous with the above, from the patient, high frequency respiratory activity signals are received such as those shown in FIG. 3(1), which signals are demodulated as appearing in FIG. 3(g) and each crossing of the zero axis by the demodulated signal is detected to `generate a series of related positive pulses as represented in FIG. 3(11), each of the pulses being shaped by a pulse shaper 41 to provide, as depicted in FIG. 3(1'), pulses 42 having sharp lead edges uniform amplitude and duration.

The pulse shaper 27 and 41 outputs, respectively representing heart beat activity and respiratory activity, are each separately fed to a box 30` (represented as 14 in FIG. l) wherein the heart beat-respiration interaction detection is carried out to detect existence of a respiratory distress condition, if present. The heart-beat pulse representations are fed into an integrator 28 which continuously integrates over an eight second time period to provide an output as illustrated in FIG. 3(d). To more vividly demonstrate what is occurring, reference is made to FIG. 4 disclosing over a longer period of time a typical EKG signal 56 that would be fed to amplifier 2S, and the lower waveform 57 represents an output heart rate waveform that would be generated by integrator 28. Note the increase in heart rate at 58 with the relative frequency increase in heart beat activity at 59. In the event, and during the period the heart rate output of integrator 28 would equal or fall below the preadjusted ninety beats per minute equivalent on the low level sensor an output signal would be generated as pulse 32 shown at FIG. 3(e) and appear as one input to AND gate 33. The second input to AND gate is a respiratory activity signal fed from pulse shaper 41 via the adjustable time delay multivibrator set at say fifteen seconds so that a signal will appear at AND gate 33 when the timefinterval between the leading edges (or even trailing edges) of successive pulse from pulse shaper 41 equals or exceeds the preadjusted period of fifteen seconds to indicate presence of a respiratory arrest, When signals at both inputs of AND gate 33 are simultaneously present the AND gate will be enabled to activate alarm 46 noting the presence of a respiratory distress condition and thus the need by the neo-natal patient for immediate attention.

Such heart rate respiration interaction detection might also be observed with reference to FIG. which shows heart rate 61 and respiration rate 62 waveforms being the respective outputs of integrators 28 and 47. It is observed that the integrated heart rate waveform 61 will fall below the ninety heart beats per minute low level between points 63 and 63, during which time a signal is present at one input to AND gate 33. It is also noted that the integrated respiration rate value 62 has a smooth discharge trend between points 64 and 64 denoting the absence of any breath during this period as otherwise an output pulse from a shaper 41 would positively alter this portion of waveform 62. t

Clearly then, subsequent to point 64 and initial point 64', there is no output pulse from pulse shaper 41, thereby indicating the presence of a respiratory arrest. Accordingly, at point 65 or exactly fifteen seconds from point 64, multivibrator 43 will be energized'to provide an output signal to appear at an input to AND gate 33, which output signal will continue to be present until point 64' when a pulse representative of a subsequent breath is emitted from pulse shaper 41 to de-energize multivibrator 43. As may be seen from FIG. 5, during the time period represented by shaded area 66 signals will appear simultaneously at the outputs of low level sensor 32 and multivibrator 43 to enable AND gate 33 denoting the presence of a respiratory distress condition as a result of the interaction detection to activate alarm 46 via OR gate 45 indicating an infant is in need of immediate attention.

As was previously noted, that in addition to the interaction detection, it was observed that in some respiratory distress cases extended respiratory arrests occurred which were not immediately accompanied by low heart rate. Very often such an extended rispiratory arrest by itself is not detectable by machine for the reason that intermittent breath action will cloud or obscure the actual physical impairment, and for this reason it has been found to be prudent to detect such extended respiratory arrests by utilizing the vehicle of respiratory rate by continuously integrating respiratory activity over a relatively small time period such as thirty seconds, and then noting how long such a low respiratory rate will continue and after the condition persists subsequent to termination of a prescribed time period, an alarm indicative of respiratory distress will be activated. Therefore, to supplement the interaction detection, circuitry including integrator 47,

low limit sensor 51 and timer delay 54 was described with reference to FIG. 2. In operation, should the output of integrator fall below a low respiration rate of, for example, eighteen breaths per minute as shown in FIG. 3(1), (integrated over a thirty second period), this low rate will be sensed by level sensor 51 which will respond by providing an output signal to trigger or start timer delay 54. Should the low respiratory rate continue for a period equal to at least sixty second timer delay 54 will be enabled, as illustrated at 55, in FIG. 3(m), to activate alarm 46 via OR gate 45. With reference to FIG. 5 waveform 62 represents the output of integrator 47, along which at point 67 the rate will fall below eighteen breaths per minute to trigger timer delay 54, however, since the rate stays below eighteen breaths per minute for twelve seconds until point 68, the low limit level sensor S1 will be de-energized and the timer delay 54, by way of lead 53, will be reset.

While preferred embodiments of the invention have been shown and described, various modifications may be made therein without departing from the spirit and scope of the invention, and it is desired, therefore, that only such limitations shall be placed on the invention as are imposed by the prior art and as set forth in the appended claims.

What is claimed is:

1. An apparatus for monitoring a respiratory distress condition through the physiological functions of heart beat and breathing comprising:

generating means for deriving an electrical signal representative of heart beat activity and an electrical signal representative of respiratory activity,

means responsive to said generating means for relating said heart beat activity electrical signal to said respiratory activity electrical signal for determining a condition of respiratory distress, and

indicator means connected to said relating means for responding to and indicating an occurrence of said condition of respiratory distress.

2. An apparatus according to claim 1 wherein said relating means includes:

timing means responsive to said respiratory activity electrical signal to detect a respiratory arrest exceeding a pre-selected time interval.

3. An apparatus according to claim 2 wherein said relating means includes:

sensor means responsive to the heart beat activity electrical signal for producing a signal upon the occurrence of successive heart beats below a preselected frequency level.

4. An apparatus according to claim 3 wherein said relating means further includes: ,t

electronic gate means responsive to coincidence of signals emitted from said timing means and said sensor means.

5. An apparatus according to claim 4 wherein said generating means includes:

a first pulse generator means adapted for connection to a patient by a suitable transducer means for deriving said electrical signal representative of heart beat activity, and

a second pulse generator means adapted for connection to the patient by a suitable transducer means for deriving said electrical signal representative of respiratory activity.

6. An apparatus according to claim 3 wherein said sensor means includes:

means for continuously integrating said heart beat activity over a constant time period of anywhere between approximately five to twenty seconds.

7. An apparatus according to claim 1 wherein said relating means includes:

sensor means responsive to the heart beat activity electrical signal for producing a signal upon the occurrence of successive heart beats Ibelow a pre-selected frequency level.

8. A monitoring apparatus for use with newborns for detecting respirator distress by investigating the physiological functions of heart -beat and breathing comprising:

first generating means adapted for deriving from a newborn electrical impulses representative of heart beat activity,

second generating means adapted for deriving from the newborn electrical impulses representative of respiratory activity,

integrating means responsive to said rst generating means for producing output signals representative of heart beat rate,

sensor means responsive to said heart beat rate to provide an output when heart beat rates are sensed below a preselected heart beat frequency, timing means responsive to said second generating means to provide an output signal when a respiratory arrest exceeds a preselected time interval,

means for relating the outputs of said sensor means and said timing means to detect a condition of respiratory distress, and

indicator means connected with said relating means for responding to and indicating an occurrence of said condition of respiratory distress.

9. A monitoring apparatus according to claim 8 wherein said indicator means includes: j.

means for providing an alarm of said condition of respiratory distress.

10. A monitoring apparatus according to claim 8 whereby:

said timing means is adjustable and may be set for detecting a respiratory arrest exceeding a time interval anywhere ybetween fifteen and forty-ve seconds.

11. A monitoring apparatus according to claim 8 wherein:

said integrating means is adapted to continuously provide a heart beat rate output representative of heart beat activity over a period of anywhere from approximately five to twenty seconds.

12. A monitoring apparatus according to claim 8 wherein:

said relating means includes an AND gate enabled upon simultaneous occurrence of an output from said sensor means and said timing means.

13. A monitoring apparatus according to claim 8 including:

second integrating means responsive to said respiratory impulses for producing output signals representative f respiration rate,

second sensor means responsive t0 said respiration rate signals to provide an output when respiration rates are sensed below a preselected respiration frequency. 14. A monitoring apparatus according to claim 13 including:

alarm means responsive to said relating means and said second sensor means. 15. A monitoring apparatus according to claim 13 wherein:

said second sensor means includes timer delay means activated by respiration rates below said preselected respiration frequency to provide an output signal subsequent to a pre-established time interval and whereby said timer delay means is deactivated and reset when respiration rates exceed said preselected respiration frequency. 16. A monitoring apparatus according to claim 15 wherein:

said second integrating means integrates said respiratory impulses over a time period of anywhere between twenty-ve to forty-five seconds. 17. A monitoring apparatus according to claim 16 wherein:

said pre-established time interval of said timer delay means is approximately sixty seconds.

References Cited UNITED STATES PATENTS 2,235,894 3/1941 Lee 12S-2.05 3,156,235 11/1964 Jaeger 12S-2.05 3,212,496 10/ 1965 Preston 12S-2.06 3,340,867 9/ 1967 Kubicek et al. 1282.05 3,347,223 10/ 1967 Pacela 12S-2.1 3,433,217 3/1969 Rieke 128-2.1X 3,438,368 4/ 1969 Karsh 12S-2.06

FOREIGN PATENTS 1,080,263 4/1960 Germany 12S-2.06

OTHER REFERENCES Lancet: Oct. 13, 1962, pp. 759-760.

Coronary Monitoring: Lexington Instr. Corp. catalogue, received in GR. 335, Jan. 3, 1968, 6 pages.

RICHARD A. GAUDET, Primary Examiner K. L. HOWELL, Assistant Examiner U.S. Cl. X.R. 12S-2.1

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U.S. Classification600/484, 600/509, 600/536
International ClassificationA61B5/08, A61B5/0245
Cooperative ClassificationA61B5/0809, A61B5/02455, A61B5/7242
European ClassificationA61B5/08J, A61B5/0245B
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Jun 28, 1982ASAssignment
Effective date: 19820526
Effective date: 19820525