|Número de publicación||US3911899 A|
|Tipo de publicación||Concesión|
|Fecha de publicación||14 Oct 1975|
|Fecha de presentación||8 Nov 1973|
|Fecha de prioridad||8 Nov 1973|
|Número de publicación||US 3911899 A, US 3911899A, US-A-3911899, US3911899 A, US3911899A|
|Inventores||Hattes Neil R|
|Cesionario original||Chemetron Corp|
|Exportar cita||BiBTeX, EndNote, RefMan|
|Citas de patentes (7), Citada por (60), Clasificaciones (7), Eventos legales (3)|
|Enlaces externos: USPTO, Cesión de USPTO, Espacenet|
United States Patent [1 1 [111 3,911,899
Hattes Oct. 14, 1975 RESPIRATION MONITORING METHOD Primary Examiner-Willam E. Kamm AND APPARATUS  Inventor: Neil R. Hattes, Danvers, Mass.
 Assignee: Chemetron Corporation, Chicago,
 Filed: Nov. 8, 1973  Appl. No.: 414,041
 US. Cl. 128/2 S; 128/l45.8; 128/DIG. 17  Int. Cl. A61B 5/00  Field of Search... 128/2 R, 2 S, 2.05 P, 2.05 R, 128/2.05 T, 2.06 A, 2.06 F, 2.06 R, 2.08, 2.1 R, 2.1 Z, 145.5, 145.8, DIG. l7, DIG. 29
 References Cited UNITED STATES PATENTS 3,316,902 5/1967 Winchel et a1 128/l45.5 3,556,083 l/l97l Grichnik et a1. 128/2.1 Z 3,575,162 4/1971 Gaardev 128/2.06 A 3,584,618 6/1971 Reinhard et a1. 128/2.1 R 3,612,041 10/1971 Ragsdale l28/2.06 A 3,690,143 9/1972 Day l28/2.08 3,768,459 10/1973 Cannon et al. 128/2 S TRANSMITT ER Altorney, Agent, or Firm.1ones, Tullar & Cooper ABSTRACT Circuitry for measuring and displaying periodically occurring physiological parameters such as the breathing rate of a patient and for providing an alarm upon the occurrence of abnormal rates is disclosed. Patient breathing is detected by variations in the distance between transmsitter and receiver coils, with respiratory movement producing a change in the magnetic field detected by the receiver coil. The resulting current is amplified and shaped to produce a breathing waveform. This signal is fed to an inspiration detector which is reflexive so that a second breath cannot be detected until there has been a definite expiration of the first breath. The output from the inspiration detector is a series of pulses representative of the breathing rate; these pulses are fed to an alarm circuit which compares the detected rates with values preset by an operator. The rate pulses are also fed to a lamp to provide a visual output indicative of the occurrence of a breath and are fed to a rate-meter for displaying the number of breaths per minute averaged over a 30 second time interval.
29 Claims, 12 Drawing Figures 1 TEST TRANS. rm MODE con. WAVEFORM 4 /|2 |s I8 22 OUT CLOCK RECEIVER ga WAVEFORM mswmmou RATE DIGITAL COlL END GENERATOR DETECTOR METER, DISPLAY 1 20 26 3o INSPIRATION men RATE VISUAL RATE LOW RATE 34 OUTPUT -36 CONTROLS AND APNEA VISUAL ALARMS AND AUOIBLE 42 OUTPUTS US. Patent ()ct. 14,1975 Sheet4 0f8 l|| mu US. Patent Oct. 14, 1975 Sheet 6 of8 3,911,899
FIG 6 NO 0 E T PRIOR ART FIG
T C E T ET DC E T E D FIG.6
L w I0. M G G S r c H W1 WI L" o 2 w 2 9 I l 2 i i m J! F l 2 Q Q Z FIG 5 LLNVERTER SHAPER I l I l I I I l OSCILLATORJ RESPIRATION MONITORING METHOD AND APPARATUS BACKGROUND OF THE INVENTION The present invention relates to detectors and'monitors, and more particularly to an improved apparatus and circuit for measuring and displaying parameters such as the breathing rate of a patient and for producing alarms upon the occurrence of selected abnormalities.
It is well known that patients undergoing intensive medical care, such as premature infants, persons suffering from respiratoryillness or persons undergoing or recovering from surgery, may often be subject to irregularities or abnormalities in breathing. These abnormalities may appear as an increase in breathing rate above a selected maximum, may appear as a gradually decreasing breathing rate which falls below a selected minimum, or may even involve an unexpected stoppage in breathing (known as apnea neonatorum). In any of these abnormal situations, the patient is in critical danger and must be given immediate assistance to prevent death or permanent damage. Accordingly, a reliable monitoring device that is capable of producing positive alarm signals upon the occurrence of selected variations from the normal breathing rate of the patient is a necessity. In addition, since the monitoring device must be capable of use with a wide variety of patients whose normal breathingrates mayv vary considerably, such a monitor must also have a wide range of adjustability to enable it to respond promptly to alarm conditions. However, the device must also be capable of differentiating between signals produced by the actual breathing motions of the patient, and signals generated by other movements of the patient or by other events that might interfere with the proper operation of the monitor.
The prior art includes numerous respiration monitoring devices which have attempted to meet the requirements of accuracy, reliability, and sensitivity required for patient safety and for convenience and ease of use by the attending operator of the unit. Such developments have related to the particular method of detecting inhalations and exhalations of the patient, methods of handling the signals resulting from such measurements, and methods of displaying the detected breathing rate. Thus, the prior art has disclosed methods of detecting the flow of air through the patients nose or the variations in the impedance of the patients chest cavity to provide anelectrical signal representative of the breathing rate. The art has also disclosed various arrangements of flashing lights and audible alarms to provide a warning of abnormalities or an indication of the ongoing breathing rate. However, such systems have not been entirely satisfactory in all applications, and efforts to develop improved'systems have continued. Such efforts have been directed toward the reduction of the complexity of prior systems and circuitry, not'only to improve reliability but to reduce the expense of such' monitors. Unfortunately, such attempts have often led not only to a simplification of the circuitry but to a corresponding reduction in the ability of the device to provide the'information required by the operator, thereby effectively reducing the usefulness of g and when an alarm condition occurs must be able to determine in the shortest period of time not only the type of abnormality occurring, but the seriousness of the problem. Thus, an alarm indicating a problem should do more then just produce a flashing light; it should also indicate that the problem is, for example, a high breathing rate and should show what thatrate is so that the operator can determine the seriousness of the problem.
SUMMARY OF THE INVENTION Accordingly, it is an object of the present invention to provide an accurate and reliable monitor and alarm system which is capable of producing audible and visual indications of patient condition.
It is another object of the present invention to provide an improved method of monitoring the respiration rate of a patient and to provide circuitry for a respiration monitor whereby the actual breathing rate of the patient is measured, converted to an electrical signal, and the resulting signal processed to provide visual and audible indications of patient condition.
It is another object of the present invention to provide an improved respiration monitor which provides visual and audible alarms at preselected breathing rates -to provide immediate and positive indications of breathing abnormalities.
Briefly, in a preferred form the subject monitor system incorporates a pair of magnetic sensors that are placed on the patient to detect patient breathing. One of the sensors is a transmitter coil which generates a magnetic field; the other sensor is a receiver coil which produces an output signal upon the occurrence of a variation in the strength of the magnetic field received. The two coils are placed on opposite sides of the chest cavity of the patient, whereby motion caused by the patients breathing varies the magnetic field strength at the receiver. A capacitor connected across the receiver coil forms a resonant circuit whereby the receiver is primarily responsive to the transmitted frequency. This signal is amplified, filtered and rectified to convert it into a varying direct current which is then further amplified and integrated before being fed to an inspiration detector. The detector is reflexive in nature so that it cannot detect a second breath until there has been a definite exhalation of the first. An automatic gain control is employed in the receiver circuit to automatically adjust the monitor for different sized patients, and the sensitivity is such that movements as small as l millimeter can be detected.
The output from the inspiration detector is a series of pulses which are representative of the breathing rate.
These rate pulses are shaped and fed to an alarm circuit which compares the detected rates with values preset by an operator. The alarm circuit incorporates a threephase clock circuit which produces three output pulses in sequence each time a rate pulse is received. The first .phase output of the clock activates all of the alarms except apnea so that any that have been previously enabled will sound. The second phase output is fed to the alarm timers to clear them for measurement of the next following time interval, and the third phase output starts the timers to initiate measurement of the next interval.
Three alarm timers are provided: one for high rate, one for low rate, and the third for detecting apnea. Each timer is adjustable by means of manual settings and after being cleared by the occurrence of a first breath start their timing operation. If the second breath occurs before the high rate timer has timed out, a high rate condition has occured and an output is applied by the high rate timer to a counter. When three consecutive high rates occur, the high rate counter provides an alarm output. If the second breath occurs after the high rate timer times out, no output is directed to the high rate counter.
When the second breath occurs after the low rate timer times out, a low rate condition has occurred, and a low rate alarm is sounded. In the preferred embodiment of the invention, the low rate alarm is reset as soon as the breathing rate is restored to the normal range.
The apnea timer detects the condition where a preset period of time elapses between successive breaths. This period is preselected by the operator and may, for example, be 10, 20 or 30 seconds. In one embodiment, the apnea timer produces two outputs, one halfway through the selected period and the other at the end of the period. In an alternate embodiment, the halfway signal is eliminated. In the first embodiment, if no breath has occurred at the halfway point, a counter register provides an output which lights a warning signal. At the end of the selected apnea period, the register provides a full alarm output unless a breath occurs. The apnea timer and the halfway signal register are reset by the occurrence of a breath.
With each succeeding inspiration pulse, the various rate timers are cycled to produce alarm conditions whenever the elapsed time between successive inspira tion pulses is greater than or less than the preselected intervals. The inspiration pulse is also fed to a ratemeter such as that disclosed in U.S. application Ser. No. 402,678 of Theodore B. Eyrick and Neil R. Hattes, filed Oct. 2, 1973 for Respiration Ratemeter, now U.S. Pat. No. 3,887,795, issued June 3, 1975, which application is assigned to the assignee of the present application, the disclosure of which is hereby incorporated by reference in the present application. As more fully described in the aforesaid application, the ratemeter displays an averaged respiration rate which is periodically updated. The inspiration pulses fed to the ratemeter are doubled by a frequency doubler and the resultant pulse train is fed in parallel to three counters. Each counter counts the input pulses for 30 seconds, after which time its count is transferred to a storage register. The counter outputs are selected in sequential order so that every 10 seconds the output from a different counter is transferred to the storage register to update the register with a new 30 second count. The storage register operates a digital optical display to provide a visual readout of the accumulated time-averaged counts which, because of the frequency doubling at the input, represents the number of breaths per minute, averaged over 30 seconds, and updated every 10 seconds.
The inspiration pulses may also be fed to a visual display which is illuminated each time the patient inspires. The system also incorporates additional features such as an alarm silencer for reducing the level of the audible alarms, and latching alarms which can be deactivated and reset only manually after the occurrence of an alarm, thus assuring that the alarm will continue until the operator takes some action.
The circuit components utilized in the system are conventional, commercially available, logic elements which are available on an off-the-shelf basis. The use of solid state elements improve the reliability and durability of the system, while allowing an improved design and better performance. Accordingly, the invention meets the requirements of simplicity and high reliability and in addition provides the operational features that have been found to be most desirable in the equipment of this type.
BRIEF DESCRIPTION OF THE DRAWING The foregoing and additional objects, features, and advantages of the present invention will be more fully appreciated from a consideration of the following detailed description of a preferred embodiment thereof, as illustrated in the accompanying drawings, in which:
FIG. 1 is a generalized block diagram of the respiration monitor of the present invention;
FIG. 2 is a more detailed block diagram of the transmitter portion of the monitor;
FIG. 3 is a more detailed block diagram of the re- FIG. 12 illustrates the relationship of FIGS. 9 and 10.
DESCRIPTION OF A PREFERRED EMBODIMENT Turning now to a more detailed consideration of a preferred embodiment of the present invention, reference is made to FIG. 1 wherein the respiration monitor is illustrated as including a transmitter coil 10 and a receiver coil 12 which respectively generate and detect a magnetic field that varies in accordance with the inspiration and expiration motions of the patient. These coils are simply a plurality of turns, for example, 988 turns of number 36 wire wrapped around a suitable form to provide an air core coil. The transmitter coil receives a high frequency alternating signal from a transmitter network 14 which may incorporate a free running oscillator operating at, for example, 4096 Hz. The magnetic field generated by the transmitter coil is picked up by the receiver coil and the resulting signal is applied to a receiver circuit turned to the transmitted frequency. In the present system, the coils l0 and 12 are secured to the patients skin by means of adhesive tape or the like on opposite sides of the patients chest, preferably at a location where motion caused by breathing will be maximized. When the patient breaths, the distance between the two coils 10 and 12 will be varied, increasing as the patient inhales and decreasing as the patient exhales, thereby producing a variable magnetic field strength at the receiver coil.
The received signal is applied to the receiver network which includes a front end circuit 16 incorporating suitable amplifiers, filters, and rectifiers to be described. After amplification and detecting, the received signal is fed to a waveform generator which filters out transients and produces an inspiration waveform signal on line which is representative of the patients breathing pattern. This waveform may, if desired, be displayed or otherwise recorded by way of line 22. The inspiration signal is applied to an inspiration detector 24 which is a high gain, reflexive, slope detector that produces a breath rate pulse on line 26 each time the patient inhales. The reflexive nature of the detector 24 requires that the patient exhale so that the inspiration signal on line 20 changes the direction of its slope before a succeeding breath rate pulse will be produced, thereby assuring that the monitor will not respond to improper breathing patterns. In addition, this feature reduces the chance that signals which might be caused by body motions unrelated to respiration will produce signals which might otherwise act as inhalation signals.
The breath rate pulse on line 26 is fed by way of line 28 to a ratemeter 30 which counts the successive pulses and produces a visual output on the digital display 32, in the manner described in the aforesaid application Ser. No. 402,678. Pulses on line 26 are also applied by way of line 34 to an inspiration visual output lamp 36 which is illuminated each time a pulse is produced on line 26. Finally, the breath rate pulses are applied by way of lines 26 and 38 to an alarm network 40 which operates to sound alarms in the event of abnormal intervals between succeeding breath pulses. The alarm network produces output signals which activate visual and audible indicators 42 in accordance with the detected breath parameters to provide the operator of the monitor with a continuing, accurate and reliable indi cation of the patients respiration.
Turning now to FIG. 2 there is illustrated a block diagram of a preferred form of the transmitter utilized with the present invention. As illustrated, the transmitter 14 may incorporate a conventional Wien bridge oscillator 44 which is free running at 4096 Hz. The output from the oscillator is applied by way of line 46 to a conventional power driver-amplifier 48 the output which is fed through a cable 50 leading to the patient and connected to the transmitter coil 10. The transmitter oscillator drives the transmitter coil which in turn produces the magnetic field that is to be detected by the receiver coil. The output of bridge 44 is also applied by way of lines 46 and 52 to a wave shaping network 54 which produces a train of pulses on line 56. These pulses may be fed to the ratemeter 30, for example, to act as its clock source.
The oscillating magnetic field transmitted by coil 10 is sensed by the receiver coil 12 and, as illustrated in FIG. 3, is fed to an automatic gain controlled amplifier 58 which is a part of the front end of the receiver portion of the monitor system. The receiver coil may be an air core coil which is connected across a capacitor to form a tank circuit resonant at the 4096 Hz transmitter frequency. The output of this tank circuit is fed by way of line 60 to the amplifier 58. The amplified signal is applied by way of line 62 through a second amplifier 64 to an active band pass filter which passes the desired signal. The output of filter 66 is again amplified in amplifier 68 and then passed through an active high pass filter 70 which has a center frequency fof lOOO Hz. The output of this filter is again amplified by an amplifier 72 and fed by way of line 74 to a detector rectifier and automatic gain control level adjusting stage 76 which converts the received signal into a direct current and controls its operating point. The pulsating DC signal is then fed by way of line 78 to a first integrator 80 which has a time constant of 125 ms. This integrator filters out transients in the received signal which may be due to patient body movements other than respiration or to other sources of noise. The output of integrator 80 is fed by way of line 82 to a buffer amplifier 84 and by way of line 86 to a second integrator 88 which has a time constant of 30 seconds. The output of integrator 88 is fed by way of line 90 to a second buffer amplifier 92.
The two buffer amplifiers 84 and 92 feed an adjustable output level differential amplifier 94 by way of lines 96 and 98, with the output of amplifier 94 producing on lines 20 and 22 an inspiration signal which corresponds to the breathing motion of the patient. The inspiration signal on line 20 is the result of a comparison in differential amplifier 94 of the output produced by the integrator 80 with the output produced by the integrator 88. The output of integrator 80 responds rapidly to changes in the signal picked up by the receiver coil, while the long term integrator 88 provides a base line for the comparison for any given coil separation. The difference between the outputs of the two integrators, then, produces the signal on line 20.
The output of buffer amplifier 92 is also applied by way of line 100 to an automatic gain control driver amplifier 102 which in turn produces the automatic gain control voltage level for amplifier 58, which is utilized at the amplifier in known manner.
The output from differential amplifier 94 is fed by line 22 to a waveform buffer amplifier 103, which may in turn feed a suitable waveform recorder or the like, while the output on line 20 from amplifier 94 is applied to the inspiration detector 24. This detector includes a high gain reflexive level detector amplifier 104 which is AC coupled to the output of amplifier 94 to allow variations in the patient waveform base line level. The output from the inspiration detector amplifier 104 is fed through a signal shaper 106, the output of which is applied by way of line 108 to an inspiration pulser 110. The pulser responds to each signal on line 108 to produce a breath rate'pulse on line 26 which is of constant amplitude and predetermined duration for use in activating the ratemeter, the visual indicators, and the alarm system.
The receiver portion of the system also incorporates a test circuit which disables the inspiration detector network and replaces the inspiration signal pulse by a two second test clock. When a test switch 112 is acti-' vated, a test latch network 114 produces an output which activates a test lamp 116 by way of line 118, inhibits the output of the inspiration detector by way of line 120, and enables a timer 122. The timer periodically activates the pulser by way of line 124 to produce an artifical 30 breath per minute inspiration rate pulse on line 26. This artificial rate pulse, which remains until the test switch is shifted from its test to its normal mode, permits the operator to test the operabil- V ,ity of the alarm and display system.
In the embodiment thus described, the breath rate pending application Ser. No. 402,677 of Theodore B. Eyrick, Allen C. Brown, and Neil R. Hattes, entitled Volume-Rate Respirator System and Method, filed Oct. 2, 1973, and assigned to the assignee of the present application. As disclosed in that application, the respirator control circuitry produces an inspiration pulse that signals the beginning of patient respiration, whether machine controlled or patient controlled. The alarm portion of the present invention which is to be described hereinbelow with respect to FIGS. 4, 9, and 10 may respond to such an inspiration pulse from a respirator to provide the herein disclosed monitoring functions in the same manner that it responds to the breath rate pulse appearing on line 26 of FIG. 3. Of course, where the alarm portion of the monitor is used with such a respirator system, the transmitter and receiver portions of the herein disclosed monitor may be ommited. Although recognizing that the input inspiration (or breath rate) pulse may be generated in either manner, in this case the alarm portion will be described in terms of its overall system receiver.
The signal on line 26 corresponding to the patient breathing rate is applied by way of lines 26 and 38 to the alarm system of FIG. 4, to which reference is now made. Generally speaking, the alarm system receives the breath rate pulses, compares the time delay between succeeding pulses to operator-selected time periods for a high breathing rate, a low breathing rate and apnea. An alarm condition occurs whenever the repetition rate of the input pulses is outside the ranges selected by the operator, and the system responds to such conditions to produce suitable visual and audible alarm signals.
The breath rate pulse is applied to a three-phase clock network 126 which responds to each input pulse to produce a series of three output signals on lines 128, 129 and 130. As will be described, the output pulse appearing as phase one ((#1) on line 128 is applied to the high and low alarm registers 132 and 134 to cause the corresponding alarms to be sounded if either of these registers has been previously enabled. The phase two (#22) signal on line 129 is applied to a reset network 136 to reset the breath interval timers, and the phase three (3) signal on line 130 may be used to restart the alarm timers.
The reset network 136 permits manual resetting of the timers by means of a reset switch 138 but also servesas a first breath lockout device which inhibits the alarms for 60 seconds after the unit is initially turned on and also inhibits the system until the first breath after a reset has occurred. The inhibition of the alarm operation for 60 seconds is accomplished by a power on clearing network 138 which responds to the initial application of power to the system to energize a not ready lamp 140 by way of line 142 and to produce an inhibiting signal on line 144 which is applied to the reset network 136. The power on clearing signal applied to line 142 is also fed through an OR gate 146 to produce on line 148 a blanking signal which inhibits the ratemeter display lamps until a meaningful breathing rate signal is present and further serves to disable the alarm system for the first minute after power has been turned on to allow the front end portion of the receiver to adjust to patient parameters.
The output of the reset network 136 is applied by way of lines 150 and 152 to the high rate register 132, by way of lines 150 and 154 to the low rate register 134 and by way of lines and 156 to the apnea register 158. It will be noted the signal on line 150 inhibits the alarm registers until after the reset and first breath lock out has been released. The reset network also produces an output of line 160 which is applied by way of line 162 to a high rate timer network 164, by way of line 166 to a low rate timer network 168 and by way of line to an apnea timer network 172. This signal clears the respective timers 164, 168 and 172 each time a breath rate pulse is received so that the timers are continually measuring the elapsed time between sucessive input pulses.
The (#3 output from clock 126 is applied by way of line 130 and line 174 to restart the high rate timer 164 after it has been cleared and is applied by way of line 176 to the low rate timer 168 to restart that timer, again after it has been cleared. Apnea timer 172 is self starting, and begins to run at the termination of the (#2 reset signal applied by way of lines 160 and 170.
The high rate timer 164 is set by an adjustable panel control such as a potentiometer 178 to time out after a predetermined interval and produce an output signal on line 180. If a breath rate pulse appears on input line 38 before timer 164 times out, the breath pulse has arrived more quickly than desired by the operator of the system (as established by the setting of potentiometer 178) and a high rate condition has occurred. This causes the l clock pulse on line 128 to shift into the high rate register 132 a data pulse indicating the exis tence of the high rate. When three consecutive high rates occur, register 132 is latched by an output appearing on line 182. The signal from line 182 is applied to a high rate lamp 184 to provide a visual indication of the condition, and is applied by way of line 186 to an input of an OR gate 188 which, in turn, produces an alarm signal on line 190. This alarm signal is fed through an audible alarm amplifier 192 to activate a suitable alarm 194.
If the high rate timer 164 times out before a succeeding breath rate pulse is received, the (b1 signal on line 128 will not insert a data pulse into register 132, the (b2 clock pulse will reset timer 164, and the 4:3 signal will restart it to measure the time interval to the next received inspiration pulse.
The low rate timer 168 measures a time interval selected by a low rate control such as a potentiometer 196 which may be adjusted by the system operator to select the lower limit for the patient breathing rate. If a measured breath occurs after the time period of timer 168 has elapsed a low rate condition is signaled on output line 198 which is applied to the low rate register 134. Thus, if the timer 168 times out before the breath is received, the d l clock pulse on line 128 will produce a data input to the register 134 which, in turn, will produce an output signal on line 200 indicative of the exis tence of an alarm condition. The signal on line 200 energizes a low rate lamp 202. In addition, the signal is applied by way of line 204 to the OR gate 188 to activate alarm 194. Since register 134 is not self latching, the alarm condition will be removed as soon as a rate of breathing above the alarm setting, as determined by potentiometer 196, is reached. Again, the Q52 and (13 clock signals clear and restart the low rate timer to permit measurement of the next breath interval. If the next breath occurs prior to the timing out of timer 168, the (bl signal to register 134 will produce no data input, and when it is reset, the alarm will not be energized, or
if it had been energized by a preceding signal, will be deenergized.
The apnea timer 172 detects the absence of any breath over a predetermined time as selected by a manually operable apnea switch 206. Since apnea is the absence of breathing, this time does not rely upon the occurrence of a next succeeding breath to trigger the alarm as do the high and low rate alarms. Accordingly, the apnea register 158 is not activated by the output from the phase clock network 126. Timer 172 is cleared each time a breath pulse is received, and then immediately begins timing the period to the next received pulse. If no pulse is received in the interval set by switch 206, timer 172 times out and produces an output on line 208 which is applied to the apnea register 158. In one embodiment of the invention, register 158 may incorporate two stages, so that the first timing out of timer 172 does not signal an alarm condition but may produce an output signal which illuminates a low rate lamp. Timer 172 is an astable, or free running, unit which at the end of its first time period resets itself to initiate a second time period to again produce a signal on line 208 when it times out. When this second signal is applied to the two stage register it produces an output signal on line 210 which is fed to an apnea indicator lamp 212 and, by way of line 214, to the OR gate 188 to energize the audible alarm 194. When the apnea alarm is sounded, the low rate indicator lamp (if used) is extinguished and the apnea register is held in the alarm condition until a breath pulse is received. Upon occurrence of a breath, the apnea time is reset by the 2 clock pulse from clock network 126.
From the foregoing, it will be seen that the present system responds to the detection of inspiration pulses to produce visual and audible alarms upon the occurrence of preset conditions which have been determined by the operator of the monitor to be dangerous to the patient. Since the desired breathing patterns of different patients and different states of health will vary widely, the system incorporates broad range adjustments to a considerable variation in the timed intervals. Thus, the high rate alarm may be set for breath rates between 20 and 100 inhalations per minute whereby rates exceeding the set limit will produce an alarm condition. Similarly, the low rate timer may be set to sound an alarm when the breathing rate falls below between 4 and 40 inhalations per minute, and the apnea timer may be set, for example, to produce an alarm condition if there is a 10, 20 or 30 second delay in which no inhalation is measured. Both visual and audible alarms are provided, with the audible alarm including silencing means for use when desired, for example, when a patient might be unduly disturbed by the sound. The system also provides a visual indication of the occurrence of each inspiration, and in addition provides a numerical readout of the number of breaths per minute as measured by the ratemeter.
A preferred form of the invention as embodied in a monitoring system which has actually been constructed and operated, is illustrated in the remaining figures, to which reference is now made. In these figures, which provide a schematic diagram of the respiration monitor, the diagrammatic blocks of FIGS. 2, 3 and 4 are indicated and similarly numbered. The circuit components utilized in this system are conventional and commercially available logic elements manufactured by numerous companies. For example, Texas Instruments TYPE DESCRIPTION 7400 Quadruple Z-Input Positive Nand 7402 Quadruple 2-Input Positive Nor 7404 Hex Inverters 7410 Triple 3-Input Positive Nand 74 I 3 Inverter-shaper 7420 Dual 4-Input Positive Nand 7474 Dual D-Type Edge-Triggered Flip-Flop Monostable Multi Vibrators In addition to the foregoing logic elements, the system also utilizes type NE555 timers manufactured by Signetics COrporation, and type LM741C Operational Amplifiers, type LM3900N Quad Operational Amplifiers and type LM370AGC amplifiers by National Semiconductor of Santa Clara, California.
Turning now to FIG. 5, there is illustrated a transmitter 14 which incorporates a Wien bridge oscillator 44 having an operational amplifier 216 connected in conventional manner to function as an oscillator, which, therefor, need not be described in detail. A variable resistor 218 provides a fine adjustment to insure proper operation at the desired 4096 Hz frequency which appears on oscillator output line 46.
The alternating high frequency signal on line 46 is applied to the base of a transistor Q1 in the power driver amplifier 48 which also acts as an impedance matching buffer amplifier. The output of the amplifier is applied by way of line 50, which may be a cable leading to the patient, to the transmitter coil 10, described above. The oscillator output on line 46 is also applied by way of line 52 to a level-shifting transistor Q2 in the wave shaper 54. The output of transistor O2 is applied to the input of a Schmitt trigger 220, the level shifter 02 being provided to insure that the oscillator signal applied to line 46 is compatible with the trigger 220. The Schmitt trigger produces on its output line 56 a train of square wave pulses having a frequency of 4096 Hz which may be used to provide a clock pulse for other monitor functions such as the ratemeter 30.
The alternating magnetic field generated by the ap plication of alternating current pulses to transmitter coil 10 is detected by the receiver coil 12, with the strength of the signal received at the coil 12 being dependent upon the distance between the transmitter and receiver coils. In addition, during inspiration and expiration of the patient, the distance between the two coils will vary so that the receiver coil will sense a changing magnetic field which has a constant frequency and is amplitude modulated by the breathing motion of the patient. The receiver coil, as indicated in FIG. 6, in-
cludes a winding 222 connected by cable 60 across a capacitor 224, the capacitor serving to tune the receiver winding to the frequency of the transmitter, thereby reducing the effects of transient or stray magnetic fields and other noise which might affect the system. The signal picked up by the coil 12 is fed by way of cable 60 to the automatic gain controlled amplifier 58 which incorporates a solid state amplifier 226 connected in conventional manner to suitable bias sources to provide a gain of about 40db. The gain of amplifier 226 is controlled by the output of the automatic gain control driver 102, received by way of line 228. Amplifier 58 is employed to permit operation of the monitor system over a very wide range of spacing between the transmitter and receiver coils, thereby permitting these coils to be located on patients having chest sizes varying from the very small, such as would be the case with a newly born infant, to large distances associated with adults whose respiration is to be monitored.
The output of AGC amplifier 58 is fed by way of line 62 to the second stage alternating current amplifier 64 which incorporates a conventional amplifier module 230 to provide an additional gain of approximately 10. The output of amplifier 230 is applied by way of line 232 to the active band pass filter 66. Filter 66 employs a pair of amplifier modules 234 and 236 which are configured in known manner to filter out undesired components of the alternating current signal appearing on line 232. The signal on line 232 is applied to an input of amplifier 234 through a series resistor 238, across a variable resistor 240, and through a series resistor 238, across a variable resistor 240, and through a series capacitor 242. The output 244 of amplifier 234 is fed back through a resistor 246 to the amplifier input and through a capacitor 248 to the junction of resistor 238 and capacitor 242. The output line 244 is also connected through the resistor 250 to an input of the amplifier module 236 with the output signal appearing on output line 252 of amplifier 236 being fed back to its input through resistor 254. In addition, the output appearing on line 252 is fed back to the junction of filter input resistors 238 and 240 by way of resistor 256, whereby variable resistor 240 is operable to provide a fine adjustment of the center frequency, which is 4096 Hz.
The filter output signal on line 252 is applied to the third stage amplifier 68 which again includes a conventional amplifier module 258 connected and biased in conventional manner to provide a gain of about 10. The output signal from amplifier module 258 is applied by way of line 260 to the high pass active filter 70 which includes an AC amplifier module 262 biased in conventional manner. The signal on 260 is applied through an RC network comprising series capacitors 264 and 266 and shunt resistor 268 to the input of amplifier 262. A feedback loop from the output 270 of amplifier 262 incorporates an RC network comprising capacitor 272 and resistor 274. Filter 70 preferably has a cutoff frequency of about 1000 Hz, thereby attenuating signals below that frequency and passing the higher signals by way of line 270 to the fourth amplifier 72.
' This amplifier again incorporates a conventional amplifier module 276 and includes in its feedback loop a capacitor 278 which serves to integrate, and thus smooth out, noise spikes appearing in the alternating current signal. I
The output of amplifier 276 is applied by way of line 280 to the detector-rectifier and automatic gain level adjustment stage 76 which converts the alternating current signal on line 280 to a DC signal and controls its magnitude. The signal on line 280 is applied through a series capacitor 282 and resistor 284 to the base of a transistor Q3. Also connected to the base of O3 is an adjustable bias source consisting of resistor 286, potentiometer 288, and resistor 290 connected in series between a bias source and ground, the slide arm of potentiometer 288 providing, through line 292, a variable bias voltage for the transistor. This potentiometer serves to adjust the operating point of transistor O3 to insure a proper range in the magnitude of the automatic gain control voltage applied to amplifier 58. The emitter of transistor 03 is connected across an RC network 294 whereby the transistor and the RC network cooperate to detect the varying amplitude of the received signal to produce on output line 78 a DC signal which varies in amplitude with the motion of the receiver coil 12 with respect to the transmitter coil 10 and which, therefore, corresponds to changes in chest dimensions of the patient during respiration.
The output signal on line 78 is fed to the first integrator 80 which comprises, as shown in H6. 7, a series resistor 296 and a parallel capacitor 298. This integrator has a time period T of 125 ms to filter out transients in the received signal caused by noise within the system, stray signals, and transients created by patient body movements other than respiration. Thus, integrator 80 serves to provide a stable DC respiratory signal which is applied by way of line 82 to a buffer amplifier 84 hav ing a unity gain operational amplifier module 300. Amplifier 300 serves as an impedance match between the integrator 80 and succeeding circuitry to prevent any drain on the capacitor 298 that might produce an erroneous respiration signal.
The output of integrator 80 is also applied by way of line 86 to the second integrator network 88 which consists of a series resistor 302 and a parallel capacitor 304. This integrator preferably has a time constant T of about 30 seconds to produce on its output line 90 a DC level corresponding to the base line around which the patient's breathing is occurring. Thus, the integrator 88 produces a reference signal which represents the average distance between the transmitter and receiver coil and it is this signal to which the relatively short term variations appearing on line 82 are compared to obtain the desired respiration signal. Because in a normal breathing cycle the patient exhales for a longer period of time than he inhales, the signal on line 90 will tend to represent the spacing of the coils after expiration, while the signal on line 82 will tend to represent the inhalation phase of the breathing cycle.
Resistor 302 of integrator 88 is bypassed by reverse connected low leakage diodes 306 and 308 which are nonconductive for any normal excursion of the respiratory signal, but which will be forward biased by a large magnitude signal such as would occur as a result of a large change in the separation of the coils l0 and 12. Such a signal will usually occur when the sensors are first being applied to a patient, or when there is a sudden large change in body position. These diodes will then conduct to bypass resistor 302 and thereby quickly change the level of the base line signal on line 90.
The output signal on line 90 is applied to the buffer amplifier 92 which incorporates a unity gain operational amplifier module 310. As with bufi'er 84, amplifier 310 provides impedance matching between the integrator and following circuitry and prevents drains on capacitor 304 that mightproduce an erroneous respiration signal.
The outputs of amplifiers 300 and 310 are applied by way of lines 96 and 98, respectively, to the differential amplifier 94 which incorporates an operational amplifier module 312. The outputs of the two buffers 84 and 92 are applied to different inputs of the differential amplifier, which then produces on its output line 20 an amplified signal portional to the difference between the outputs of buffers 84 and 92, which outputs are at the same level as the voltages appearing across capacitors 298 and 304, respectively.
If the output of integrator 80 were used alone for the production of breath rate pulses, no usable signal would be available, for the base line of the integrator would be different for each different separation of the sensor coils. With the high gain of amplifier 312, its output would be saturated most of the time and a usable pulse would not be available on the differential amplifier output line 20. The integrator 88 acts to integrate the changes appearing across capacitor 298 and thus is used as a reference for those changes. The amplifier 312 serves to amplify the difference between the outputs of the two integrators, not the absolute value of either, and accordingly the output signal on line 20 will be proportional to the difference in the relative position. of the coils 10 and 12 after inhalation and after exhalation.
Because inhalation separates the transmitter and receiver coils and produces a weakening signal, inhalation produces a decreasing value across capacitor 298 of integrator 80; however, it is desired to represent that voltage as an increasing value in order to provide the desired inhalation pulse, and for this reason the output of buffer 84 is applied to an inverting input of amplifier 312, in the preferred embodiment. It will be noted that amplifier 312 is provided with an adjustable bias by means of potentiometer 314 which permits adjustment of the output of amplifier 312 to a quiescent level of about volts.
The output of buffer amplifier 92 is also applied by way of line 100 to the automatic gain control driver 102. The signal on line 100 corresponds to the base line of the respiratory signal and is fed to the base of transistor O in the driver 102. The output of transistor O5 is fed to the base of a second driver transistor Q6 which controls the gain of the automatic gain control amplifier 58, as has been noted. Thus, the time period of integrator 88 becomes the time constant for the AGC amplifier 58, allowing the unit to self adjust its gain for use on any size patient but at the same time taking a sufficiently long time (30 seconds) to do so to insure that changes caused by the breathing of the patient will not be attenuated by the change in gain of amplifier 58.
The respiration waveform appearing at the output of amplifier 312 is applied by way of line 322 to the waveform driver 103 which incorporates a transistor amplifier 04 connected in grounded emitter configuration to act as a low output impedance. The output of driver 103 may be applied by way of output line 318 to an oscilloscope, to a pen recorder, to a volume-flowrate calculator, or the like, as desired.
The output signal from the differential amplifier appearing on line is a waveform representative of the actual breathing profile of the patient, and is fed to the inspiration detector 104 which operates to produce an output pulse for each inhalation by the patient. The inspiration detector includes an operational amplifier 320 configured as a high gain comparator having a positive feedback resistor 322 and an offset adjustment potentiometer 324. The input of amplifier 320 is AC coupled to the output of the differential amplifiers by way of capacitor 326, and because of this AC coupling, am-
plifier 320 acts as a slope detector rather than as an absolute level detector. Because of the positive feedback through resistor 322, the output of amplifier 320 will change only with a change in the slope of the input waveform, rather than a return to a specific predetermined level.
The value of a slope detector in this circuit is that it can detect patient breathing during a changing base line, whereas a circuit that requires a return to a preset level could not detect such a breathing pattern. This is illustrated in graphs A and B of FIG. 8, with graph A representing a level detector and graph B representing the slope detection of the present system. In graph A, the breathing waveform 328 is shown as returning to a base line between each detection of a breath. In such a system, the inspiration detector turns on when the waveform increases above a set level, and turns off when the waveform returns to a preset lower level. If the base line is changing, however, as when the patient takes a series of inhalations with only a partial exhalation between each, there can be no detection of the subsequent inhalations with such an arrangement. In graph B, on the other hand, where the waveform 328' goes through a series of level changes, each without returning to the original base line, the present system is still capable of producing an output pulse for each new inhalation, as long as it is preceded'by sufficient exhalation to cause a reversal of the slope of the breathing waveform.
It will be noted that the level of input signal which turns on amplifier 320 is not adjustable, but such adjustability is not required in the present system since the amplifier 320 is not responsive to values but to changes in slope. Accordingly, it will be seen that amplifier 320 is turned on when the input waveform on line 20 changes in a positive direction and is turned off when the waveform changes in a negative direction, thereby producing at its output line 330 a pulse having a repetition rate corresponding to the inhalation rate of the patient. The output signal on line 330 is limited by a diode 332 to prevent it from going below 0 volts in order to protect the following stage of the receiver system, and is applied to a pulse shaper 106. The pulse shaper consists of a Schmitt trigger 334 which acts to provide a square wave pulse on line 108 in response to each input pulse, except when inhibited by a test signal on line 118, fed to the trigger by way of line 336. This inhibiting signal prevents pulses produced by the inspiration detector from being fed to the remainder of the system during a test mode.
The square wave pulse appearing on line 108 is fed to one input of an OR gate 338 in the inspiration pulser 110, the output of gate 338 being fed through an AND gate 340 to a one shot multivibrator 342 to produce output signals on lines 26 and 28 which have uniform durations. The multivibrator preferably has a duration of 0.3 seconds and its output pulse is connected to the ratemeter 30 by way of line 28 and to the inspiration lamp driver 344 and lamp 346 by way of lines 26 and 34. The lamp is illuminated each time a pulse occurs, while the ratemeter accumulates and displays the accumulated pulses in the manner set forth in the abovereferenced US. Pat. No. 3,887,795. The signal on line 26 is the breath rate pulse which corresponds to the rate of inspiration of the patient and which is applied by way of line 38 to the alarm circuitry of the monitor illustrated in block diagram form in FIG. 4 and to be described in detail heeinbelow.
To test the operation of the monitor, the test mode switch 112 is provided which, when shifted to the test position enables the latching circuit 114 formed by an OR gate 348 and an AND gate 350. The resulting signal on output line 118 activates test lamp 116, comprising a driver 352 and a lamp 354, to provide an indication that the system is in the test mode. The signal is also applied through line 336 to disable the pulse shaper 106, as has been described, and is applied through an inverter 356 to one input of an AND gate 358, enabling this latter gate. The test circuit incorporates timer 122 which comprises a free running multivibrator 360 that produces an output pulse on its output line 124 every 2 seconds. This timer pulse is applied to the other input of AND gate 358 so that when the gate is enabled by the test mode signal, the timer pulses appear on line 362 for application to the OR gate 348 in inspiration pulser 110. These timer pulses are then used to activate the one shot multivibrator 342 to produce internally generated inspiration pulses on lines 26 and 28. The timer pulses activate the ratemeter to display a 30 breaths per minute rate and may also be used to check the operation of the alarm circuits by setting the high rate and low rate timers to 30 breaths per minute. If the sensor coils and 12 are not attached to a patient, apnea can be checked by turning off the test switch, thereby removing the internally generated breath rate pulses and causing the apnea alarm to sound after the preset time.
The inspiration pulses on line 38 are fed to the three phase clock network 126 (FIG. 9) which responds to each inspiration pulse to provide pulses in sequence on lines 128, 219 and 130. The l pulse on line 128 strobes the alarm circuits to produce an alarm signal if the high or low rate alarms have been enabled, the 2 signal on line 129 clears and resets the alarm timers, and the phase three signal on line 130 restarts the various timers, in the manner to be described.
The three phase clock network 126 incorporates a four-bit shift register having flipflops 364, 365, 366 and 367. A timer 368 provides clock pulses for each of the flipflops, the timer being free running to produce a train of pulses having an on time of 6 ms and an off time of 0.2 ms. This train of pulses is applied by way of line 370 to the clock inputs of flipflops 364-367 by way of lines 372, 373, 374 and 375, respectively. The inspiration pulse on line 38 is fed through an inverter 376 and by way of line 378 to the drive input of flipflop 364.
In the presence of an inspiration pulse at the drive input of flipflop 364, the timer output pulses on line 370 are able to shift the flipflops 364-367 sequentially to produce the three output phases required for operation of the alarm. Thus, the first timer pulse which occurs after the inspiration pulse is received will shift flipflop 364 to its set condition, causing its output line 380 to go high, and its output line 382 to shift low. The high signal on line 380 is applied to the data input of flipflop 365 so that upon generation of a second output pulse by timer 368, flipflop 365 will be set. Again, the setting of flipflop 365 will produce a high signal on line 384 and a low signal on line 386 whereby the third pulse from timer 368 will set flipflop 366. The setting of the latter flipflop causes it to shift, producing a high signal on its output line 388 and a low signal on its output line 390. Finally, the high signal on line 383 is then applied to flipflop 367 so that the fourth timer output pulse will set flipflop 367, shifting its output 392 to a low value.
The three phase clock circuit 126 incorporates three AND gates 394, 396 and 398 which are interconnected with the various flipflops and with the timer 368 to produce the three output phase pulses on lines 128, 129 and 130. The inputs to AND gate 394 are connected to output lines 380 and 386 of flipflops 364 and 365, respectively, and to the timer output line 370 by way of line 400. When flipflop 364 has been set, so that line 380 is high, and before flipflop 364 has been set, so that line 386 is high, AND gate 394 will produce an output signal on line 370 also to go high. The output pulse on line 128 will last as long as all three inputs to AND gate 394 are high so that the resulting l signal on line 128 has the same duration as the timer pulse. When flipflop 365 sets, its output line 386 goes low so that the AND gate 394 will not conduct when the next timer pulse is received. However, after flipflop 365 sets and before flipflop 366 sets, AND gate 396 is enabled by the high signals on lines 384 and 390. Accordingly, the next clock pulse from timer 368 will be applied to AND gate by way of line 402 so that all of the inputs to gate 396 will be high, thereby producing the (#2 pulse on line 129. At the end of the clock pulse on line 402, the signal on line 129 terminates.
When flipflop 366 sets, AND gate 396 will no longer be conductive, and the high signal on line 388 is applied to one input of AND gate 398. As long as flipflop 367 is still reset, a high signal appears on its output 392 and a high output from timer 368 applied by way of line 404 to AND gate 398 will produce the (#3 signal on line 130. It will be noted that flipflops 364-367 remain set as long as the inspiration pulse is present on line 38 so that only one series of phase pulses will be applied to lines 128, 129 and 130 for each input pulse on line 38. At the end of the inspiration pulse, flipflops 364-367 are clocked to their reset condition by timer 368 in preparation for the next inspiration pulse.
When the monitor is first turned on, some time is required for the system to stabilize and for the various indicators to reach values that will correspond to the breathing rate of the patient being monitored. In order, therefore, to prevent incorrect indications or premature sounding of alanns, the power on network 138 is provided to disable the alarms and to provide a blanking signal to the various display lamps for a period of time sufficient to enable the system to reach stable operation. The power on circuit includes a timer 406 which is controlled by a power sensing network comprising resistor 408, capacitor 410 and diode. 412. When power is first applied to the system, and biasing voltages are applied to the timer 406, the power sensing network presents a low level input to timer 406 by way of line 414, triggering the timer which then produces an output pulse on its output line 416. As capacitor 410 charges, the input on line 414 rises so that timer 406 is freed to time out after 1 minute. The resulting power on pulse on line 416 is applied through AND gate 146 and line 418 to a display blanking network 420 to disable the various digital display lamps used in the monitor for the duration of the time delay imposed by timer 406. A diode 422 serves to isolate the segments of one display lamp from those of another lamp.
Diode 412 is provided across resistor 408 to provide a fast discharge bypass for capacitor 410 when power is turned off so that if power to the monitor should be turned off and then back on again almost immediately, the level on capacitor 410 will be sufficiently low to restart timer 406.
A lamp test switch 424 is gated through AND gate 146 by way of line 426, closure of the switch disabling the AND gate to permit unblanking of the display lamp segments if it should be desired to test the lamps during the first minute of operation of the monitor. If desired, the signal on line 416 may be applied through a line 428 to a driver amplilfier 430 to operate the test indicator lamp 116.
The output of timer 406 is applied by way of line 144 to the reset and first breath lockout network 136 to disable the system alarms for the one minute delay imposed by timer 406. Network 136 further prevents operation of the alarms until a first breath is detected, after which the system operates in the normal manner. The signal on line 144 is applied to one input of an OR gate 432 in reset network 136, with the output of OR gate 432 being applied to the drive input of a flipflop 434. The output of this OR gate is also applied to the direct reset terminal of the flipflop by way of line 436, thereby producing a high signal on output line 438 which is fed through OR gate 440 to produce a resetting signal on line 160. As explained with respect to FIG. 4, the signal on line 160 resets all of the system timers as well as the apnea register, thereby disabling the alarm system. The alarm remains disabled until after the power on timer 406 times out and the first breath is received.
The system may also be manually reset by means of reset switch 138 which, when closed, provides a signal by way of line 442 through an inverter 444 to a second input of the OR gate 432, whereby closure of switch 138 direct resets flipflop 434 to disable the alarms and to reset the timers. Finally, a reset signal may be applied to line 160 by the occurrence of a (#2 signal on line 129 which is applied through an inverter 446 to a second input of OR gate 440.
At the end of the power on delay period, the signal on line 144 is removed, and, if the reset switch 138 is open, flipflop 434 is no longer held in direct set by the output of OR gate 432. The next inhalation pulse that appears on line 38 will, therefore, produce a clock pulse on line 129 which will be fed through inverter 446 to OR gate 440 and through line 448 to the clock input of flipflop 434, setting the flipflop to free the system alarm timers. A high signal is also produced on output line 150 from flipflop 434 which resets the alarm registers 152, 134 and 158. Thereafter, a (#3 signal on line 130 initiates the breath rate timers by way of lines 174 and 176.
The 3 signal is applied to the high rate timer 164 (FIG. by way oflines 130 and 174. This signal is fed to a monostable multivibrator timer 450 which has a variable duration of from 0.5 to 2.4 seconds, depending upon the setting of a control potentiometer 452 in the high rate control 178. When timer 450 times out, it produces a low output signal on line 454 which is applied to one input of an OR gate 456 where the signal is-inverted, and thence to an inverter 458, which then produces a low output. This signal is applied by way of line 180 to the data input of a first flipflop 460 inthe high rate shift register 132. i
The receipt of the next inspiration pulse by clock network 126 after the starting of ,timer 450 will produce a 1 signal on line 128 which is applied to the clock input of flipflop 460 through an inverter 462. If this next inspiration pulse has occurred after timer 450 has timed out, then the clock input flipflop 460 is low, leaving the flipflop in its reset condition, which is not an alarm condition. However, if timer 450 has not timed out when the new breath pulse is received, then the data input to flipflop 460 will be high when the (1:1 signal on line 128 is received and accordingly this flipflop will be clock set. When this occurs, the output line 464 of flipflop 460 will go high, freeing the direct clear input of the succeeding flipflops 466 and 468 in the shift register 132. The line 464 is connected to the data input of flipflop 466, so that if the next succeeding breath pulse again occurs before the timer has timed out, the (#1 signal will be applied through inverter 462 to the clock input of flipflop 466 by way of line 470, setting flipflop 466. Since the data input to flipflop 460 will still be high, that will also remain set. Similarly, if the third successive inhalation pulse is received before timer 450 times out, flipflop 468 will be set, producing a highsignal on output line 182 indicative of an alarm condition.
It will be understood that timer 450 is reset by the Q52 signal after each successive inhalation pulse, so that it measures the time that has elapsed between each of the successive pulses. If the breath interval is less than the selected interval, as determined by potentiometer 452, for each of three consecutive breaths, an alarm signal is produced on line 182. This alarm signal is fed back through OR gate 456 by way of line'472 to latch the shift register in its alarm condition. The signal on line 182 is also applied by way of line 186 through first and second OR gates 474 and 476 of the OR network 188 and appears on line 190 for application to the alarm driver 192. Driver 192 comprises a pair of transister amplifiers Q7 and Q8 which activate a suitable alarm device 194 such as the commercially available Sonalert manufactured by the Mallory Company.
If desired, the alarm device may be provided with an alarm volume selector 478 having a selector switch 480 by means of which the alarm may be silenced, may be caused to sound at a reduced level, or may be operated at its maximum volume. Preferably, the silient position of the switch is connected through an indicator lamp 482 to provide a visual indication of the alarm setting. It will be further noted that the output of the OR network 188 may be connected by way of line 190 through transistor switch O9 to an alarm jack 483 to permit the monitor to drive external alarm devices.
If a breath pulse interval greater than the length of time set by timer 450 occurs before flipflop 468 is set,
' thereby indicating a nonalarm breath interval before three consecutive alarm intervals are measured, then flipflop 460 will be clock reset, producing a low signal on its output 464 which direct resets flipflops 466 and 468. Thus, the only way that a high rate alarm signal can be present on line 182 is through the occurrence of three consecutive high rate breaths. When this occurs, the alarm is sounded, shift register 132 is latched A single interval outside the preset limit will cause the alarm to sound, but a subsequent non-alarm breath interval will turn the alarm off. As before, the first breath following the termination of the power on delay produces a (#3 starting signal on line 130 which is applied by way of line 176 to a monostable multivibrator timer 484 in the low rate timer 168 (FIG. At the end of the time period set by a potentiometer 486 in the low rate timer selector 196, a low output signal appears on timer output line 488 which is applied through an inverter 490 to the low rate timer output line 198. The signal on line 198 is fed to the data input of a flipflop 492 in the low rate register 134. If the next succeeding breath after initiation of timer 484 occurs before the timer has timed out, the data input to flipflop 492 is low, it is left in its reset condition, and there is no alarm. However, if timer 484 has timed out, its output will be low, the output of the inverter will be high and, when the (b1 pulse on line 128 caused by the next succeeding breath pulse appears and is fed through an AND gate 494 to the clock input of flipflop 492, that flipflop will be set. This produces an output on line 200 which is fed through an inverter 496 to energize the low rate lamp 202 and applied a signal by way of line 204 to the second input of OR gate 474 in the OR network 188 (FIG. 9), thereby driving the audible alarm 194.
Whether or not the timer 484 has timed out, the (#2 signal from the next following inhalation pulse will clear the timer by way of lines 160 and 166 from the reset network 136 and the timer will be restarted by the 3 signal on line 176. Since the state of the timer output is clocked into flipflop 492, the low rate alarm can be reset as well as set, so that the alarm is activated only as long as an alarm condition actually exists. If the breath succeeding an alarm condition breath is received within a nonalarm interval, the low rate flipflop will be reset and the alarm silenced. The AND gate 494 provides the one exception to this resetting operation. If an apnea condition has been detected, but not yet reset, AND gate 494 will be disabled to prevent the resetting of the low rate alarm; by the (b1 clock pulse on line 128, even if the low rate timer has been reset.
The apnea timer 172 differs from the preceeding timers in that it does not rely upon a measurement of the interval between successive breaths, since apnea is the absence of breathing. Instead, timer 172 operates to produce an alarm after a given length of time has elapsed since the occurrence of the preceeding breath. Accordingly, the apnea timer network 172 incorporates an astable, or free running multivibrator timer network 498 which has its time period controlled by a three position apnea switch 206. Switch 206 incorporates a pair of movable switch contacts 500 which are manually adjustable to select one of three values of resistance, whereby, in one embodiment, the timer 498 will operate at 5, 10 or second intervals. These timer periods may be one half those of the desired apnea intervals so that two timer cycles are required to produce an apnea alarm, in which case a two stage apnea register 158 is provided. It will be understood, however, that the selector switch may be used to select the full apnea interval, in which event a single stage flipflop register may be used to signal apnea.
In the present description, a two-stage apnea alarm is shown, and thus two cycles of the timer 498 are required to produce an alarm. When timer 498 times out, its output on line 502 goes low. This signal is applied through and AND gate 504 to the timer output line 208, which feeds it to the clock input of a first flipflop 506 in the apnea register 158. Output line 508 of flipflop 506 is connected to the clock input of the following flipflop 510 and also to the data input of flipflop 506 whereby the occurrence of a first time out signal on line 208 will clock set flipflop 506. Since timer 498 is astable, when it times out the first time, it will begin a second cycle. If no new breath occurs during this second on time, the timer will again time out to clock reset flipflop 506 and, in turn, clock set flipflop 510. The latter flipflop provides the apnea signal which is fed by way of line 210 to the apnea indicator lamp 212 and by way of line 214 to the second input of OR gate 476 in the OR network 188 (FIG. 9) to energize the alarm. The setting of flipflop 510 also produces a high signal on output line 512 of flipflop 510 which is fed back to the second input of AND gate 504 to hold the output of that AND gate low and prevent further shifting of register 158. The signal on line 512 is also applied by way of line 514 to disable AND gate 494 to prevent the low rate register from setting while there is an apnea alarm.
If a breath pulse occurs before the apnea register goes to its alarm condition, a reset pulse will be applied by way of lines 60 and 170 to reset timer 498 and by way of lines and 516 to reset flipflop 506, thereby preventing an apnea alarm condition.
It will be noted that the output line 382 of flipflop 364 in the three phase clock network 126 (FIG. 9) is connected to the data input of flipflop 510. This is done to prevent the flipflop 510 from setting if it receives a clock pulse from flipflop 506 which is the result of a #12 signal resetting of flipflop 506, thereby preventing a false alarm condition.
Thus there has been disclosed a respiration monitor which is capable of detecting patient breathing by sensing variations in the distance between two magnetic sensors that are placed on the patient. One sensor generates a magnetic field and the other receives any variations of the signal strength caused by the breathing motion of the patient. An automatic gain control system is employed to automatically adjust the monitor for different size patients and the system operates to produce inspiration pulses for each inhalation motion of the patients chest occuring after an exhalation motion. The system is sensitive to the direction of motion rather than the absolute value in order to provide an acurate monitoring of the actual breathing function. The receiver portion of the system produces inhalation pulses which are applied to a ratemeter to provide a digital display averaged over a 30 second time interval. Every 10 seconds a new respiration rate is computed, presenting the average breathing rate for the previous 30 seconds. The inspiration pulses also are applied to an alarm system which measures the breath to breath intervals to monitor high breathing rates or low breathing rates monitors the occurrence of apnea, and provides alarms and illumination of appropriate indicators in alarm conditions. The high rate alarm triggers after a delay of three breaths to compensate for any noise signals, such as may be caused by movements of the body other than respiratory movements, while the low rate alarm triggers immediately upon detection of an excessive time period between breaths. The apnea alarm operates when no breath is received for a predetermined selected period of time. An inspiration indicator flashes each time the patient inspires, and the audible alarm may be adjusted for sound intensity levels or may be silenced completely. Once a high rate or apnea alarm has been triggered, it can only be deactivated and reset by manual operation of the alarm reset switch. Although the present invention has been described in terms of a preferred embodiment used for measuring patient respiration, it will be apparent that the inventive concepts described herein may be applied to the monitoring of other periodic physiological parameters. Suitable measuring equipment may be used to detect such parameters and to produce pulses similar to the inspiration pulses produced by the herein-described receiver, and the alarm system of the present invention may then be used to provide a warning of abnormal pulse repetition rates signalling alarm conditions. Thus, it will be seen that numerous changes and modifications can be made by those of ordinary skill on the art without departing from the true spirit of scope of the invention as defined in the following claims.
What is claimed is:
1. In a monitor, circuitry for measuring the rate of recurrence of a physiological parameter of a patient and for providing an alarm upon variation of the parameter from preselected values;
sensing means responsive to said parameter for producing a sensing signal; first integrator means responsive to said sensing signal to produce a variable parameter signal;
second integrator means responsive to said variable parameter signal to produce a relatively stable base signal; differential amplifier means responsive to the difference between said base and variable parameter signals to produce an electrical signal having a waveform the slope and amplitude of which varies in accordance with the parameter being monitored;
reflexive means responsive to said electrical signal for detecting said rate of recurrence, said reflexive means being responsive to a recurring slope of said waveform, whereby a waveform having a slope of a first direction initiates a pulse, said reflexive means preventing the production of a second pulse until said waveform reverses to a second direction and returns to said first direction to thereby define a complete cycle of said parameter and a recurrence of the parameter being measured;
first adjustable alarm means responsive to the rate of recurrence of said pulses to produce a first alarm signal when said rate exceeds a first preset value;
second adjustable alarm means responsive to the rate of recurrence of said pulses to produce a second alarm signal when said rate falls below a second preset value; and
indicator means responsive to said alarm signals.
2. The monitor of claim 1, wherein:
said first integrator means has a relatively short time constant; and
said second integrator means has a relatively long time constant.
3. The monitor of claim 2, wherein said sensing means further includes a sensor responsive to magnetic fields modulated in accordance with variations in said parameter to produce said electrical signal, said parameter being the inspiration and expiration motions of the patient, whereby said waveform of said electrical signal corresponds to the breathing profile of the patient.
4. The monitor of claim 1 wherein said first alarm means incorporates a first alarm timer for measuring intervals between successive pulses and producing an excessive pulse rate output signal whenever the interval between two successive pulses is less than a predetermined period, and first register means responsive only to the occurrence of three successive excess pulse rate output signals from said first timer to produce said first alarm signal.
5. The monitor of claim 4, wherein said second alarm means incorporates a second alarm timer for measuring intervals between successive pulses and producing a low pulse rate signal whenever the interval between two successive'pulses is greater than a predetermined period, and second register means responsive to a low pulse rate signal from said second timer to produce said second alarm signal.
6. The monitor of claim 5, further including third adjustable alarm means responsive to the rate of recurrence of said pulses to produce a third alarm signal if a following pulse is not received within a preset period of time after the end of a prior pulse.
7. The monitor of claim 6, wherein saidthird alarm means incorporates a third alarm timer, and third register means responsive to an output signal from said third timer to provide said third alarm signal.
8. The monitor of claim 7, further including reset means for resetting said third alarm timer upon the occurrence of a pulse prior to the completion of the preset period of time for said third timer.
9. The monitor of claim 1, wherein said reflexive means comprises amplifier means having an input and an output;
positive feedback means connected between the output and the input of said amplifier; and
AC coupling means coupling said electrical signal to the input of said amplifier means whereby said amplifier is responsive to changes in the slope of the waveform of said electrical signal to produce at its output pulses having a repetition rate corresponding to the rate of recurrence of said parameter.
10. In a monitor, circuitry for measuring the rate of recurrence of a physiological parameter of a patient and for providing an alarm upon variation of the parameter from preselected values, comprising:
sensing means for producing an electrical signal having a waveform which varies in accordance with said parameter;
detecting means responsive to the waveform of said electrical signal to produce rate pulses corresponding to the occurrence of said parameter; first adjustable alarm means responsive to the rate of recurrence of said rate pulses to produce a first alarm signal when the rate of recurrence exceeds a first preset alarm condition value;
second adjustable alarm means responsive to the rate of recurrence of said rate pulses to produce a second alarm signal when the rate of recurrence falls below a second preset alarm condition value;
third adjustable alarm means responsive to the rate of recurrence of said pulses to produce a third alarm signal if a following pulse is not received within a preset period of time after the end of a prior pulse;
a clock pulse generator responsive to each said rate pulse to produce first and second phase pulses;
|Patente citada||Fecha de presentación||Fecha de publicación||Solicitante||Título|
|US3316902 *||25 Mar 1963||2 May 1967||Tri Tech||Monitoring system for respiratory devices|
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|Clasificación de EE.UU.||600/407, 128/205.23, 600/595|
|Clasificación internacional||A61B5/113, A61B5/11|
|25 Sep 1992||AS||Assignment|
Owner name: ALLIED HEALTHCARE PRODUCTS, INC.
Free format text: RELEASED BY SECURED PARTY;ASSIGNOR:BANE BOSTON FINANCIAL COMPANY;REEL/FRAME:006329/0348
Effective date: 19920831
|26 Ago 1985||AS||Assignment|
Owner name: BANCBOSTON FINANCIAL COMPANY (THE LENDER)
Free format text: SECURITY INTEREST;ASSIGNOR:ALLIED HEALTHCARE PRODUCTS, INC. A CORP. OF DE.;REEL/FRAME:004444/0863
Effective date: 19850612
|27 Mar 1981||AS||Assignment|
Owner name: ALLIED HEALTHCARE PRODUCTS, INC.
Free format text: CHANGE OF NAME;ASSIGNOR:CHEMETRON-MEDICAL PRODUCTS, INC.;REEL/FRAME:003925/0807
Effective date: 19810227