U.S. patent number 3,911,899 [Application Number 05/414,041] was granted by the patent office on 1975-10-14 for respiration monitoring method and apparatus.
This patent grant is currently assigned to Chemetron Corporation. Invention is credited to Neil R. Hattes.
United States Patent |
3,911,899 |
Hattes |
October 14, 1975 |
Respiration monitoring method and apparatus
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.
Inventors: |
Hattes; Neil R. (Danvers,
MA) |
Assignee: |
Chemetron Corporation (Chicago,
IL)
|
Family
ID: |
23639712 |
Appl.
No.: |
05/414,041 |
Filed: |
November 8, 1973 |
Current U.S.
Class: |
600/407;
128/205.23; 600/595 |
Current CPC
Class: |
A61B
5/113 (20130101) |
Current International
Class: |
A61B
5/11 (20060101); A61B 5/113 (20060101); A61B
005/00 () |
Field of
Search: |
;128/2R,2S,2.5P,2.5R,2.5T,2.6A,2.6F,2.6R,2.08,2.1R,2.1Z,145.5 |
References Cited
[Referenced By]
U.S. Patent Documents
Primary Examiner: Kamm; Willam E.
Attorney, Agent or Firm: Jones, Tullar & Cooper
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 said third 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;
means to apply each of said first phase pulses to said first and
second alarm means to produce alarm signals when an alarm condition
exists; and
means to apply each of said second phase pulses to said first,
second and third alarm means to prevent the production of alarm
signals when no alarm condition exists.
11. The monitor of claim 10, wherein said means to apply said
second phase pulses to said alarm means includes a reset network,
said reset network further including manual reset means and first
breath lockout means.
12. The monitor of claim 10, wherein said first alarm means
incorporates a first alarm timer for measuring intervals between
successive rate pulses and producing an alarm condition output
signal whenever the interval between two successive pulses is less
than a predetermined period, and first register means responsive to
a plurality of successive alarm condition output signals to produce
said first alarm signal.
13. The monitor of claim 12, wherein said second alarm means
incorporates a second alarm timer for measuring intervals between
successive rate pulses and producing a low pulse rate signal
whenever the interval between two successive rate pulses is greater
than a predetermined period, and second register means responsive
to a low pulse rate signal to produce said second alarm signal.
14. The monitor of claim 10, wherein said first alarm means
incorporates a first alarm timer for measuring the intervals
between successive rate pulses and first register means responsive
to said first timer to produce said first alarm signal upon the
occurrence of an excessive pulse rate; said second alarm means
incorporates a second alarm timer for measuring the intervals
between successive rate pulses and second register means responsive
to said second timer to produce said second alarm signal upon the
occurrence of an excessive interval between two successive rate
pulses to produce said second alarm signal; and said third 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.
15. The monitor of claim 14, further including means for applying
said first phase pulses to said first and second registers to
enable said registers to produce an alarm signal; and
a reset network responsive to said second phase pulses to reset
said first, second, and third alarm timers, whereby if an alarm
condition exists in said first or second registers, said first
phase pulse will cause an alarm signal, and if no alarm condition
exists said timers will be reset to monitor the next pulse rate
interval.
16. The monitor of claim 15, further including indicator means
responsive to said alarm signals and power on clearing means for
blanking said indicator means to prevent an indication of an alarm
condition for a predetermined time after said monitor is initially
turned on.
17. The monitor of claim 16, further including test means for
generating pulses at a preselected rate for application to said
reset means and for inhibiting the output of said detecting
means.
18. The monitor of claim 10, wherein said parameter is the
respiration rate of a patient, said sensing means comprising means
for measuring the inspiration and expiration motions of the patient
to produce a breathing waveform, and wherein said first and second
alarm means produce alarm signals in response to rates of
inspiration corresponding to said preset alarm condition values,
and said third alarm means produces an alarm signal upon the
occurrence of apnea.
19. The monitor of claim 18, wherein said sensing means comprises
first and second sensor coils and oscillator means connected to
said first sensor coil to produce an alternating magnetic field,
said second sensor coil being responsive to said magnetic field to
produce an output having an amplitude proportional to the spacing
between said sensor coils whereby relative motion of said sensor
coils amplitude modulates the output of said second sensor coil,
and receiver means responsive to said amplitude modulated output to
produce said electrical signal having a waveform which varies in
accordance with the breathing pattern of said patient.
20. The monitor of claim 10, wherein said sensing means
includes:
means responsive to said parameter for producing a sensing
signal;
first integrator means having a relatively short time constant
responsive to said sensing signal to produce a variable parameter
signal;
second integrator means having a relatively long time constant
responsive to said variable parameter signal to produce a
relatively stable base signal; and
differential amplifier means responsive to the difference between
said base and variable parameter signals to produce said electrical
signal having a waveform which varies in accordance with the
parameter being monitored.
21. The monitor of claim 20, wherein said detecting means comprises
a positive feedback amplifier AC coupled to said differential
amplifier means, said detecting means being responsive to the slope
of the waveform of said electrical signal to produce said rate
pulses.
22. A method of monitoring the respiration rate of a patient
comprising:
generating a magnetic field;
modulating said field in accordance with the breathing pattern of
the patient;
sensing said field and producing an alternating current signal
having an amplitude modulation proportional to said breathing
pattern;
converting said alternating current signal to a pulsating direct
current signal proportional to said breathing pattern;
integrating said direct current signal over a relatively short time
period to produce a first integrated signal;
integrating said first integrated signal over a relatively long
time period to produce a second integrated signal;
differentially amplifying said first and second integrated signal
to produce a breathing waveform having a base dependent upon the
value of said second integrated signal;
producing breathing pulses in response to changes in the direction
of slope of said breathing waveform;
sensing the intervals between successive breathing pulses;
comparing the sensed intervals to preselected high rate low rate
and apnea values; and
producing alarm signals when said breathing pulses occur at
intervals outside said preselected values.
23. The method of claim 22, wherein the steps of sensing and
comparing intervals and producing alarm signals comprises:
activating an adjustable high rate timer, an adjustable low rate
timer and an adjustable apnea timer upon the occurence of a first
breathing pulse;
determining whether either said high rate timer or said low rate
timer has timed out upon the occurrence of a second breathing
pulse;
producing a first alarm signal if said high rate timer has not
timed out, thereby indicating a high breathing rate;
producing a second alarm signal if said low rate timer has timed
out, thereby indicating a low breathing rate;
producing a third alarm signal if said apnea timer times out before
a second breathing pulse occurs, thereby indicating an apnea
condition; and
resetting all of said timers upon receipt of said second pulse.
24. A monitor for providing an alarm upon variation of a measured
physiological parameter from preselected values, comprising:
means producing rate pulses corresponding to the rate of recurrence
of a measured physiological parameter;
first alarm means including a first alarm timer for measuring the
intervals between successive rate pulses and producing an excess
pulse rate output signal when any interval is less than a first
predetermined period;
first register means responsive to said first timer to produce a
first alarm signal upon the occurrence of excess pulse rate output
signals corresponding to a plurality of consecutive rate pulses
each occurring at intervals less than said first predetermined
period;
second alarm means including a second alarm timer for measuring the
intervals between successive rate pulses and producing a low rate
output signal when any interval is greater than a second
predetermined period;
second register means responsive to said second timer to produce a
second alarm signal upon the occurrence of a low rate output
signal;
reset means responsive to each said rate pulse to produce a
plurality of sequential phase pulses;
means applying the first-occurring one of said phase pulses to said
first and second register means to enable said registers to produce
an alarm signal; and
means responsive to the second-occurring one of said phase pulses
for resetting said first and second alarm timers, whereby if an
alarm condition exists in said first or second registers said first
phase pulse will cause an alarm signal, and if no alarm condition
exists said timers will be reset to monitor the time period of the
next succeeding rate pulse.
25. The monitor of claim 24, further including third alarm means
including a third alarm timer for measuring elapsed time between
successive rate pulses and producing a third alarm signal if the
elapsed time between signals is greater than a third predetermined
period, said means responsive to the second-occurring one of said
phase pulses further resetting said third alarm timer to prevent
production of said third alarm signal and to initiate measurement
of the elapsed time to the next succeeding rate pulse.
26. The monitor of claim 25, further including means responsive to
a third-occurring one of said phase pulses for restarting said
first and second alarm timers.
27. The monitor of claim 24, wherein said first register is
responsive only to three consecutive excess rate pulses to produce
said first alarm signal.
28. The monitor of claim 24, wherein said parameter is the
respiration rate of a patient, and wherein said means for producing
rate pulses includes means generating breathing pulses
corresponding to the rate of recurrence of successive breath
inhalations separated by breath exhalations.
29. 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 for producing an electrical signal having a waveform
the slope and amplitude of which varies in accordance with said
parameter;
reflexive means for detecting said rate of recurrence, including
amplifier means having an input and an output;
positive feedback means connected between the output and the input
of said amplifier;
Ac coupling means coupling said electrical signal to the input of
said amplifier means, said amplifier being 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 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.
Description
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 respiratory
illness 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
breathing rates may 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
patient's nose or the variations in the impedance of the patient's
chest cavity to provide an electrical 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 the
monitor. Other attempts to provide suitable monitoring equipment
have failed to recognize the actual needs of the monitor operating
personnel, who must be able to determine the condition of the
patient quickly, 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 that rate 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 patient's
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 1
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 three-phase 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 inspiration 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 receiver portion of
the monitor;
FIG. 4 is a more detailed block diagram of the alarm portion of the
monitor;
FIG. 5 is a schematic diagram of a preferred form of the system
transmitter;
FIGS. 6 and 7 are a schematic diagram of the receiver of FIG.
3;
FIG. 8 is a graphical illustration of the operation of a
inspiration detector in the receiver;
FIGS. 9 and 10 are a schematic diagram of the alarm of FIG. 4;
FIG. 11 illustrates the relationship of FIGS. 6 and 7; and
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 10 and
12 are secured to the patient's skin by means of adhesive tape or
the like on opposite sides of the patient's 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 20 which is representative of the patient's 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 indication of the patient's
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 f of 1000 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 activated, 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
110 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 operability of the
alarm and display system.
In the embodiment thus described, the breath rate pulse appearing
on line 26 is a shaped pulse of predetermined amplitude and
duration which is produced each time the patient inhales. This
pulse is produced by the receiver portion of the monitor
illustrated in FIG. 3, in one form of the monitor system, where the
actual respiration motion of a patient is detected. In an alternate
form of the monitor, this pulse may be produced by a respirator of
the type disclosed and claimed in copending 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 (.phi.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 (.phi.2) signal on line 129
is applied to a reset network 136 to reset the breath interval
timers, and the phase three (.phi.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 serves as 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 150 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 .phi.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
.phi.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 .phi.1 clock
pulse on line 128 to shift into the high rate register 132 a data
pulse indicating the existence 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 .phi.1 signal on line 128 will not
insert a data pulse into register 132, the .phi.2 clock pulse will
reset timer 164, and the .phi.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 .phi.1
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 existence 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
.phi.2 and .phi.3 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 .phi.1
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 .phi.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 Corporation of Dallas, Texas, Signetics
Corporation of Sunnyvale California, Stewart-Warner Company of
Sunnyvale, California, and numerous other manufacturers provide the
necessary components on an off-the-shelf basis. Essentially, the
present system utilizes conventional series 7400 logic as
follows:
TYPE DESCRIPTION ______________________________________ 7400
Quadruple 2-Input Positive Nand 7402 Quadruple 2-Input Positive Nor
7404 Hex Inverters 7410 Triple 3-Input Positive Nand 7413
Inverter-shaper 7420 Dual 4-Input Positive Nand 7474 Dual D-Type
Edge-Triggered Flip-Flop 74121 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 Q2 is applied to the input of a Schmitt
trigger 220, the level shifter Q2 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 application 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, includes 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.
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 Q3 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 Q3 to insure a
proper range in the magnitude of the automatic gain control voltage
applied to amplifier 58. The emitter of transistor Q3 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 FIG. 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 having 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 10 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 buffer 84, amplifier 310 provides impedance matching
between the integrator and following circuitry and prevents drains
on capacitor 304 that might produce 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 0 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 Q5 in the driver 102. The output of
transistor Q5 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 Q4 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
20 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, amplifier 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 above-referenced
U.S. 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 10 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 .phi.1
pulse on line 128 strobes the alarm circuits to produce an alarm
signal if the high or low rate alarms have been enabled, the .phi.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 .phi.1 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 .phi.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 .phi.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 alarms, 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 .phi.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 .phi.3
signal on line 130 initiates the breath rate timers by way of lines
174 and 176.
The .phi.3 signal is applied to the high rate timer 164 (FIG. 10)
by way of lines 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 in the high rate shift register 132.
The receipt of the next inspiration pulse by clock network 126
after the starting of timer 450 will produce a .phi.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 .phi.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 .phi.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 high signal on output line 182 indicative of an alarm
condition.
It will be understood that timer 450 is reset by the .phi.2 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 Q9 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 into the alarm condition so that no
following non alarm breaths can reset it, and the system remains in
the alarm condition until it is manually reset by means of switch
138.
The low rate timer 168 compares the interval between two
consecutive breaths against the time period set by the low rate
control 196. If the breath interval is greater than the selected
interval, then the alarm is set. 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 .phi.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. 10). 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
.phi.1 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 .phi.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 .phi.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 .phi.1 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
15 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 130 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 160 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 .phi.2 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 patient's 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.
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