CA1210821A - Patient monitor for providing respiration and electrocardiogram signals - Google Patents

Abstract

PATIENT MONITOR FOR PROVIDING RESPIRATION AND ELECTROCARDIOGRAM
SIGNALS
ABSTRACT OF THE DISCLOSURE

The patient monitor includes a patient unit having a probe connected to receive a carrier signal, which probe is adapted for connection to the body of a patient to be monitored. The carrier signal is passed through the patient's body and modulated in accordance with the respirations of the patient to produce a modulated carrier signal. A carrier detection circuit is connected to receive the modulated carrier signal and produce a demodulated respiration signal. An ECG
circuit also receives the carrier signal and filters out ECG signals. The respiration and ECG signals are passed to an analysis unit. Both the patient unit and analysis unit contain baseline correction circuits for maintaining a predetermined baseline.

Description

PAT:IENT ~IONITOR FOR PR~NII)ING RESPIR~lqON AND

BACRGROUND OF TE[E INVENTION

F i eld o~ the Inve n~ i on This invention relates to systems for monitoring specific patient parameters and more parti-ularly to systems which monitoY electrocardiogram tlECG) waveforms and respiration waveformi, and which are designed to provide oultput signals which have a controlled ampli tude and baseline .

D i scus s i on o~ Rela ted_Ar t Monitoring of specific pa~ient parameters on a continuing basis is ~ecoming a generally accepte~d diagnostic tool. This is part.icularly true in the case of infants which aEe deemed "at risk" and susceptible ~q~

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'o s~dden infant death syndrome. Such infants exhibit prolonged apnea and bradycardia episodes. ~pnea is defined as ~he cessation of respiration, and brady-cardia is defined as low heart rate A presently available monitor is a Model 16000 Infant Monito ~manufactured and sold by ~eal.hdyne, Inc., of ~arietta, Gerorgia. ~his infant monitor is designed to manage infants who have been determined to be at risk by providing signals indicative of the infant's respiration and heart activity. The monitor contains two control adjust-ments which must be made by the operator to properly set ~p the unit. These controls are for the sensitivity setting of the respiration and ECG
lS channels. The monitor provides excellent operation when the sensitivity settîngs are ~roper. However, it is possible for people to poorly adjust the sensitivity settings a~d in so doing cause signal~dropouts and accompanying false alarms. Ac~ordingly, a need has 20 developed for a monitor which automatically controls the sensitivity of the respiration and ECG signals.
The present invention can be used in combination with a recorder .o provide a visual display of the monitored parameters.

SUMMARY OF T~E INVENTION

One object of the present invention is to provide a patient monitor which produces respiration and ECG output signals that are indicative of the patient's respiration and hear~ ac~ivitv, respectively.

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Another object of the present invention is to provide a patient monitor which includes a patient connected unit which receives respiration and ECG
signals from a patient, and a signal analysis unit ~hich receives the respiration and ECG signals from the patient connected unit, but in which the patient connected unit and signal analysis unit are electri-cally isolated from each other in order to eliminate the possibility of electrical shock to the patient due to malfunctioning of the signal analysis unit.
Another object of the present invention is to provide a patlent monitor having a patient connected unit and a signal analysis unit in which the patient connected unit maintains a proper baseline for received signals, and in which the signal analysis unit can control the baseline correction ~unction of the patient connected unit.
Another object of the present invention is to provide a patient monitor which has a patient connected unit which is capable of sensing the existence of a loose lead on the patient and eliminating an ou~put from the patient connec~ed unit during the presence of a loose lead.
A further object of the present invention is to provide a patient monitor in which signals are transmitted from a patient connected unit to a signal analysis unit in a manner which eliminates the possibility of signal distortion due to interference by spurious signals.
Yet another object of the present invention is to provide a patient monitor having a signal analysis unit which can detect a deviation of a signal from a proper baseline at several points in the unit.

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A further object of the present invention is to provide a patient monitor which includes an ECG
channel in which the gain of the ECG signal is controlled automatically.
A still further object of the present invention is to provide a patient monitor having a signal analysis unik in which the baseline of a signal can be restored automatically at a controlled ra~e.
In accordance with the above and other objects, the patient monitor of the present invention comprises a patient unit having a probe connected to receive a carrier signal, which probe is adapted for connection to the body of a patient to be monitored, whereby the carrier signal is passed through the patient's body and modulated in accordance with respirations of the patient to produce a modulated carrier signal. A car~ier detection circuit is connected to receive the modulated carrier signal and produce a demodulated respiration signal. An amplifier amplifies the demsdulated respiration signal and the resultant am~lified respiration signal pwlse width modulate PWM an oscillator ~o produce a PWM respiration signal, The patient monitor also includes an analysis unit which contains a carrier generation circuit for producing the carrier signal. The analysis unit also contains a pulse width demodulation circuit which receives the PWM respiration signal, demodulates the si~nal, and thereby produces a respiration data signalu An output circuit of the ~ 5 analysis unit is connected to receive ~he respiration data signal, amplify and level shift the respiration data signal and output the resultant respi.ration data signal.
The monitor also includes isola~ion cir~uitry for electrically isolating the patient unit from the analysis unit.
The patient unit also includes a baseline correction circuit for sensing the DC level of the demodulated respiration signal and adding or sub~racting a DC signal to the demodulated respiration signal in response to the sensed DC level. The base line correction circuit includes a capacitor which is charged in response to the sensed DC level. The patient monitor also includes a circuit fQr changing the charging rate o~ the capacitor when the DC level o~ the respiration data signal reache~ a predetermined amount. This charging rate circui~ comprises a circuit contained in the analysis unit for deactivating the carrier generation circuit to stop the production of the carrier signal~ and a circuit in the patient unit for sensing the lack of carrier signal and reducing the charging time constant of the ca~citor in response thereto.
The analysis unit also includes a baseline correction circuit for sensing the DC level of the res~iration data signal and adding or subtracting a DC
signal to the respiration data signal in response to the sensed DC level. The analysis unit baseline correction circuit includes a capacitor which is char~ed in accordance with the sensed DC level and includes circuitry for varying the rate of charging of the capacitor. The circuit for varying the rate of charging the capacitor includes a circuit for reducing the charging ~ime constant of the capacitor when the s DC level of the respiration data signal is above or below predetermin2d limits. This rapid charge circuitry includes a programmed microprocessor, and a pair of comparator circuits connected to receive the respiration data signal and produce outputs when the respiration data signal is above or below upper anci lower limits, respectively.
The patient monitor al50 includes an ECG
sensing circuit which includes a filter contained in the patient unit for passing frequencies associated with an ECG signal. The ECG circuit also includes an amplifier ~or amplifying the frequencies passed by the ~ilter to produce an amplified ECG signal, and a frequency modulation circuit conneeted to frequency modulate the amplified ECG signal to produce a PWM ECG
signal. The ampiifier and frequency modulation circuit are contained in the patient unit, and a frequency demodulation circuit is contained in the analysis unit for demodulating the PWM ECG signal to produce an ECG data signal. An ECG output circuit is contained in the analysis unit for receivin~ the ECG
data signal~ amplifying, level shifting and outputting that signal.
The ~CG output circuit includes an automatic gain control circuit for controlling the amplitude of the ECG data signal to within predetermined limits.
The automatic gain control circuit includes a gain controllable amplifier in the form of an operational amplifier with a variable resistance optical coupler : .

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contained in a feedback loo~, and a gain control circuit in the form of an integrator circuit having an input connected to the output of the gain controllable amplifier and having an output connected to the control input of the optical cou~ler.
The monitor also incl~des an ECG baseline correction circuit in the patient unit for sensing the DC level of the amplified ECG signal and adding or subtracting a DC signal to the amplified EGG signal in response to this sensed DC level. The ~CG baseline correc~ion circuit includes a capacitor which is charged ln accordance with the sensed DC level.
The charging time constant of capacitor of the ECG baseline circuit varied in a manner similar to the variation of the charging time constant of the capacitor in the respiration baseline correction circuit of the patient unit. That is, when the DC
level o~ the respiration data signal reaches a predetermined level, the carrier generation circuit is 2C deactivated thus sto~ping the generation of the carrier signal. A circuit in the patient unit senses the cessation of the carrier signal and reduces the charging time constants of the capacitors in both the ECG baseline correctlon circuit and the respiration-baseline correcton circuit. The carrier generation circuit is similarly deactivated due to a high DC
level of the ECG data signal.

BRIEF DESCRIPTION OF T~E DRAWINGS
. _ _ The above and other objects of the invention will become more readily apparent when the invention is more fully described below, reference being had to the accompanying drawings in which like reference numerals repcesent like par~s throughout, and in which:

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. Figure 1 is a block diagram of the patient monitor of the ~resent invention;
; Figure 2 is a schematic diagram showing the patient unit of the present invention;
Figure 3 is a schematic diagram showing the respiration channel of the signal analysis unit of the present invention;
Figure 4 is a schema~ic diagram showing the rCG channel of the signal analysis unit of the present invention~
Figure 5 is a schem~tic diagram of the carrier signal generation unit of the present invention;
Figures 6a and 6b show a flow diagram of the main program used in the microcomputer of the oresent invention;
Figures 7a, 7b.and 7c show a flow diagram of the interrupt service routine used in the microcomputer of the present inventioni 2~ Figure 8 is a schematic diayram of a power supply unit for use with the present invention; and Figure 9 is a schematic diagram of a test signal generation circuit for use with the present invention; this figure appears on the sheet of drawings containing Pigure 5.

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DETAILED DESCRIPTION OF TH PR~FERREI) .,MBODIMENT

Figure 1 shows a block diagram of the patient monitor 10 of the present invention. Patient monitor 10 is connected to a patient 12 for ~he purpose of providing electrocardiogram signals indiGative of the patient's heart activity nd resplration signals indicative of ~he patient's respiration. The signals provided by the pa~ient monitor 10 can be used to actuate alarms indicating bradycardia or apnea episodes, can be displayed on an convenient deyice such as a C~T, and can be recorded.
for future reference to assist a physician to make a diagnosis of the patient.
Monitor 10 includes a patient un~ut 14 which contains all patient connected leads. Si.gnals developed in patient unit 14 a.re transmitted to signal analysis unit 16 which performs control and analysis functions for the ~onitor. Also, analysis unit 16 contains the monitor power supply. In order to guard against electrical shock to patient 1~, low voltage levels are used in patient unit 14 and patient unit 14 is electrica~ly isolated from analysis UAit 16 SO that circuit malfunctions in .the analysis unit will not result in electrical shocks to the patient.
Carrier generator 18 is located in the analysis unit and produces a high frequency carrier signal. This carrier signal can be approximately 100 khz. This carrier signal i5 passed to patient unit 1 through an isolation transformer 20~ In patient 30 unit 14, the voltage and urrent levels of the carrier signal are adjusted in input circuit 22 and passed to patient 12 through standard car~on or silver chloride electrodes. The carrier signal is modulated in a known manner according ~o an increase or decrease in the chest expansion of the patient so as to produce an amplitude modulated signal in which the envelope indicates res~iration of the patien~. Also, the normal voltages associated with heart activity of the patient are added to this signal. The resultant signal is passed through line 24 to the respiration channel of the monitor, and through line 25 to the ECG
channel of the monitor.
Input circuit 22 also includes power supply circuitry which receives the carrier signal and rectifies and filters this signal to ~roduce a low lS level bias voltage for the patient unit 14.
The respiration channel includes a synchro-nou5 detector ~ which receives the carrier signal on line 27 and synchronously detects the amplitude modulated carrier signal received on line 24. The detected signal is low pass filtered in filter 28 and the demodulated respiration signal is passed to amplifier 30~ The demodulated respiration signal is amplified in amplifier 30 and the amplified respira-tion si~nal is passed to modulator 32~ Mo~ulator 32 is a pulse width modulator in the ~orm o~ a voltage controlled oscillator which converts the amplified ~espiration signal into a pulse width modulated respira-tion signal. This pulse width modulated respiration signal is passed to isolation device 40 and then to 3n demodulation circuit 42 of the analysis unit 16. The signal is pulse width modulated so that variations in amplitude of the signal due to, or example, spurious noise, 60 hz signals from the power supply, etc., will not affect the respiration signal adversely. The output of demodulator 42 is a respiration data signal which is passed to amplifier 44. From amplifier 44, the signal is passed to logarithmic amplifier ~8 in which the dynamic range of the signal is compressed and passed through switch A6 to output line 47 which can be connected to a recorder, CRT display device, or the like. A non-compressed, linear output is taken from amplifier 44 through switch 45 and provided on output line 51 for the same purpose. The output of amplifier 44 is also ~assed to signal sha~ing circuit 49 which can be a zero crossing detector. The output of signal shaping circuit 49 is passed through line 57 to microprocessor 56, and can al50 be passed to an end lS use device such as a strip chart recorder or cathode ray tube display. Microprocessor 56 receives the signals on line 57 and can be programmed to perform standard operations such as actuating an alarm in the event of an apnea episode.
Patient unit 14 also includes a loose lead detector 38 which senses the output of synchronous detector 26. If no out~ut is sensed from synchronous detector 26, this is interpreted as indicating that a lead has pulled loose from patient 12 and loose lead detector 38 passes an inhibit signal to modulator 32 to deactivate the modula~or so that no signal will be passed to analysis unit 16. Analysis unit 16 also contains a loose 12 d detector 50 which senses the output from isolation circuit 40. If no output is sensed, loose lead detector 50 sends a signal through line 55 to micro~rocessor 56 which may sound an audible alarm, illuminate a visible alarm, or the like.

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Patient unit 14 also contains a baseline correction circuit 36 which senses the DC level of the output of amplifier 30 and either adds or subtracts a DC si~nal via amplifier 30 to the respiration si~nal depending on the sensed level. If the DC offset of the res~iration signal due to, for example, large signal levels due to defibrillation of the patient, or the like, reaches a limit which cannot be quickly corrected by baseline correction circuit 36, the larger DC level is passed through isolation device 40 and demodulator 42 and detected by baseline detection circuit 52. Baseline detection circuit 52 sends a ~signal through line 53 to microprocessor ;6 which, in response to this signal, stops the operation of carrier generator 18 through line 60. Patient U;lit 1~l has a level detecting circuit 34 which is connected to the output of isolation transformer 20. When the carrier signal is stopped by microprocessor i6, level detector circuit 34 no longer sees an output signal from isolation transformer 20 and reduces the time constant of baseline correction circuit 36 in response. Circuit 36 then rapidly corrects the base-line deviation.
Analysis unit 16 contains a basellne correction circuit 58 which receives the outp~t of logarithmic amplifier 48 on line 59. Baseline correction circuit 58 senses the DC level of the logarithmically converted signal and adds or subtracts a DC signal to the respiration data signal in amplifier 44 based on the sensed DC level of the logarithmically converted signal. If the DC offset of the respiration data signal becomes too ~reat for base line correction circuit 58 to compensate for, the DC

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level will reach a threshold set by baseline detection circuit 54 which receives the output of amplifier 44.
Baseline detection circuit 54 sends a signal to microprocessor 56 which can con~rol the baseline correction circuit action through line 61 and fast correction circuit 63 or line 62 and slow correction circuit 64. Fas~ and slow correc~ion circuits 63 and 64 control the rate at which baseline correction circuit 58 follows the output of loqarithmic amplifier 48. Also, microprocessor ~6 can stop the carrier generation, if necessary.
The ECG channel comprises a filter 70 which passes signals in the frequency range associated with ECG signals and filters out the carrier signal. The filtered ECG.si~nal is amplified in amplifier 72 and passed to pulse width modulation circult 74. The amplified ECG signal is modulated in circuit 74 and passed through line 73 to isolation circuit 76 and demodulation circuit 7~ of the analysis uni~ 16. The ECG si~nal iS pulse width modulated in order to avoid contamination from noise. The signal is demodulated in circuit 78 and amplified in amplifier 80. The amplified signal is passed to automatic gain control circuit 84 and through switch 82 to output line 100.
Automatic gain control circuit 84 maintains the ampli-tude of the ECG data signal ~i~hin predetermined limits and passes the signal to shaping circuit 86.
Shaping circuit 86 contains a f$1~er to pass the R
wave portion of the ECG signal to microprocessor ~6 through line 87. If microprocessor 56 sees that the signal on line 87 is missing, it injects current through line 65 to automatic gain c~ntrol circuit 84 to rapidly increase the gain of that circuit to its maximum in an attempt to restore a proper signal level.

The gain control signal is also passed through line 77 and switch 101 to line 102.
The ECG channel of patient unit 14 contains a baseline correction circuit 75 for the ECG signal.
Baseline correction circuit 75 senses the output of amplifier 72 and adds or subtracts a DC signal to the ECG signal to compensate for DC drift. If the DC
drift becomes too ~reat for baseline correction circuit 75 to rapidly correct, the excessive DC level is passed through isoLation circuit 76, demodulation circuit 78 and amplifier 80. Baseline detection circuit 88 receives the signal from amplifier 80 and senses that the DC level of the signal is above or below predetermined limits. Baseline detection circuit 88 5end5 a signal through line 89 to micropro-cessor 56 which in turn ~eactivates carrier generator 18. This causes leve} detection circuit 34 to reduce the time constant of baseline correction circuit 75, which then rapidly corrects the baseline deviation.
The analysis unit 16 contains a baseline correction circuit 90 which also receives the output of amplifier 80 and adjusts the baseline of the ECG
signal by subtracting or adding a DC signal to the ECG
data signal at amplifier 80 in response to the DC
level of the ECG data signal. If baseline correction circuit 90 cannot compensate for the DC level of the ECG data signal, baseline detection circuit 88 will note an excessive baseline deviatlon and notify microprocessor 56 which, through line 62 and fast baseline circuit 93, reduces the tlme constant of baseline correction circuit 90.
In operation, carrier generator 18 passes a 100 khz carrier through isolation transformer 20 to patient unit 14. This signal is received in input circuit~22 which rectifies the signal to provide _ l5 ~

biasing voltages for the circuitry of the patient unit. Input circuit 22 a].so adjusts the voltage level of the carrier signal and controls ~he carrier si~nal current before passing the carrier signal to patient 1~. The carrier signal is amplitude modulat~ in accordance with respirations of patien~ 12 and passed through line 24 to synchronous detector 26. The amplitude modulated respiration signal is detected and filtered in circuit 28. The filtered signal is amplified in amplifier 30 which also corrects the base line of the signal so that the DC level remains at zero volts. The signal from amplifier 30 is modulated by frequency modulation circuit 32. In the event that no signal is emitted from synchronous detector 26, loose lead detector 38 disables modulator 32.
Accordingly~ modulator circuit 32 does not transmit a signal through isolation circuit 40, and loose lead detector 54 therefore signals microprocessor 56 of the existence of a lo~se lead. Microprocessor 56 can then provide an appropriate alarm signal.
Assuming that a proper signal is passed through synchronous detector 26, if the DC level of the signal is too high or too lowt this DC level is corrected by baseline correction circuit 36 in amplifier 30. Consequently, the signal passed to modulation circuit 32 should have a zero volt ~C
offset and should be amplified so that contamination due to noise will be minimized. This signal is p~se width modulated prior to being passed to the s.ignal analysis unit 16, thereby fu~ther reducing the ~ b~

possibility of noise contamination. The signal is demodulated in circuit 42. Baseline correction circuit 36 is quite effective in compensating for offset voltages due to capacïtive absorption or capa-citive voltage buildu~ due to minor perturbations inthe received signal. However, if the capacitive ~uildup becomes excessive due to, for example, defibrillation of the patient, the baseline must be corrected more rapidly. Accordingly, baseline detec-tion circuit 52 senses the excessive baseline andsignals microprocessor 56 to turn off carrier genera-tor 18. Thus, the carrier wave input to patient unit 14 is removed. As a consequence, detector 34 senses the cessation of the carrier wave and reduces the time con~tant of baseline correction circuit 36, which rapidly corrects the baseline.
At the same ~ime, the ECG voltage from the patient is passed along line 25 through filter 70 to amplifier 72 where it is amplified and passed to frequency modulator 74. Baseline correction circuit 75 maintains a proper baseline for the ECG signal.
When the carrier wave is removed. Level detection circuit 34 reduces the time constant of baseline correc~ion circuit 75. Consequently, each time that ~5 the respira~ion channel baseline is restored, the baseline for the ECG channel is also restored.
Returning to the respiration channel, it will be seen that the noise-free demodulated signal is amplified in amplifier 44 which pa~ses the amplified signal to log amp 48. The baseline of the signal in . .
., ~ 17 -amplifier 44 is corrected by baseline correction circuit 58 and an excessive deviation of the baseline is sensed by baseline detection circuit 54. Two baseline detection circuits are used so that baseline drift in any portion of the circuit will be detected and immedia,ely corrected. It can be seen that detec-tion circuit 52 will respond to any baseline drift in the respiration channel of the patient unit.
Detection circuit 54 responds to baseline drift in amPlifier 44. If the drift in the patient unit is in an opposite direction to the drift in amplifier 44, baseline detection circuit 54 will not produce an output~ Nevertheless, detection circuit 52 will produce an output thus correcting the baseline. If the drift is only caused by amplifier 44, detection circuit 52 will not produce an output but detection circuit 54 will, thus ensuring that the deviation is promptly corrected for.
Baseline correction circuit 58 normally operates in a very slow mode to follow the DC offset of the slow respiration data signal. However, when an excesslve baseline deviation is to be compensated for, circuit 58 is operated in ~he fast mode by actuating circuit 63 through microprocessor 56. In the fast operational mode, circuit 58 will rapidly follow the DC offset of the respiration data signal.
To avoid discontinuity in the baseline due to switched baseline time constants, a medium variable circuit 64 is actuated after fast correction circuit 63 i5 deactuated to allow baseline correction circuit 58 to follow the DC offset of the respiration data signal more accurately. The variable circuit can change the time constant to slow gradually over a period of several minutes following deactuation of circuit 58 Returning again to the ECG channel, it will be seen that the noise free demodulated ECG data signal is amplified at amplifier 80 and passed through switch 82 and ~GC circuit 8~ to signal shaping circuit 86. Baseline correction circuit 90 compensates for baseline drift of the ECG data signal. When the base-line drift becomes excessive, detector ~8 signals microprocessor 56 to operate the baseline circuit in a fast mode through fast circuit 93. In the system of the present invention, baseline correction circuit 90 and baseline correction circuit 58 are both operated in the fast mode at the same time, and switches 45, 46 and 82 are opened at that time so that no output is transmitted to a display device while rapid baseline .l5 correction is being e~fected.
The ECG data signal passed through AGC
circuit 84 is output to shaping circuit 86 and to output line 77 through switch 101 with an amplitude which is maintained within predetermined limits. The ECG data signal passed through amplifier 80 is outpu~
to line 100 through switch 82, but with an amplitude which is not controlled, representing 1 v output on line 100 for 1 mv input ECG signal. Shaping circuit ~6 contains a filter which is designed to pass the R
wave portion of the ECG signal. Microprocessor 56 senses the R wave portion of the ECG signal on line 87 and, if the signal is missing, causes a rapid increase in the gain of AGC circuit 84 through line h5 to ensure that the missing signal is not due to poor gain

3~ in the circuit. This feature is of much value in avoiding false bradycardia and cardiac arrest alarms after some artifact causes radical AGC gain reduction.

Figure 1 also shows a power supply 95 for the circuits of the present invention. Power supply 95 will be discussed in detail hereafter. An apnea time delay switch 96 is also shown. Switch 96 can have two or more settings and provides signals to microprocessor 5~ relating to a minimum time period for defining an apnea event. Finally, a reset switch 97 is shown. Switch 97 includes two contacts which send a ground signal, respectively, thro~gh line 98 iO to power supply 95 and through line 99 to micro-processor 56. As will be discussed in detail, in order to turn off power supply 95 or to change the position of apnea switch 96, reset switch 97 must be actuated, otherwise an alarm signal is produced~ This 15 ~eature preven~s tampering with the switch set~ings o~
the present invention.
Figure 2 shows the patient unit 14 of the present invention connected to isolation transformer 20. The primary of isolation transformer 20 is 20 connected to a 100 khz voltage source so that the secondary of the transformer delivers a 100 khz signal to input circuit 22 which includes rectifier bridge 200 that produces a full wave rectified ~C output that is used for the bias voltages of the patient unit.
25 The DC output signal of rectiier bridge 200 is filtered by capacitor~ 201 and 207~ The input circuit also includes back-to-back constant current diodes 202 and 203 and steering diodes 204 and 205. Diode 204 passes the positive half waves of.the carrier signal 30 to constant current diode 203, and steering diode 205 passes the negative half waves of the carrier signal to constant current diode 202. The output of this diode network is connected to patient 12 through a carbon or sllver chloride electrode, as is standard in 35 the art. The patient is also connected to a ~loating ground.

As discussed above, ~he ~arrier signal is modulated in the patientls body in accordance with ; res~irations so that an amplitude modulated respira-tion signal is ?roduced. ~his signal is passed along line 24 to synchronous detector 26 which comprises FET
206, the gate of which is connected thro~gh diode 210 to the floating ground and through capacitor 208 to the carrier signal input. The input to capacitor 20B
is also connected thro~gh back-to back zener diodes 212. Diode 210 clamps the gate of FET 206 to the floating ground and zener diodes 212 act as input protection for the FET. Capacitor 208 AC co~ples the carrier signal to the FET gate so that FET 206 is turned on in response to peaks of the carrier signal 15 thus providing synchronous detection for the amplitude modulated respiration signal. The detected respira-tion signal is passed to filter section 28 which AC
couples the signal through capacitor 214 to filter 218~ Clamping diodes 216 clamp ~he positive and 20 negative going portions of the signal to the floating ground to protect the ciscuit from voltages due to, for example, defibrillation of the patient. Filter section 218 is a low pass filter which smooths the signal and passes it to amplifier 220. Amplifier 220 25 contains a feedback network 222 having resistor 221 and capacitor 223. Amplifier 220 provides a substan-ti~al gain increase to the si~nal and passes it to frequency modulator 32. Frequency modulator 32 is a voltage controlled oscillator ~VCO) comprising an 3Q integrated circuit volt~e controlled oscillator 224 which can be an Intersi~ ICM 7555 integrated circuit having an apDropriate biasing network. Also, voltage controlled oscillator 24 has an enable input connected to line 241. Line 241 is connected through diode 242 35 to comparator 240. The non-.inve~ting lead of comDarator 240 is connec~ed to ehe output of~FET 206, I~
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and the inverting input connected to a positive voltage source. When a loose lead connection exists and the output of FET ri.ses to a level indicative of 150C ohms or greater, comparator 240 outputs a signal through diode 242 to disable VCO 224.
Baseline correction circuit 3~ essentially comprises a capacitor 238 which senses the Eed back output of amplifier 220 thorugh resistors 221 and 239 and charges in accordance with the DC level thereof.
The voltage develope~ on capacitor 238 is fed to the inverting input of amplifier 220 and, accordingly, the DC level at the inverting input of amplifier 2~0 equals the DC level of the non-inverting input of the ampli~ier so that ~he output has a baseline at zero volts. Baseline correction circuit 36 also includes FET 234 which shunts resistors 221 and 239. The gate of FET 234 is connected to level detection circuit 34.
Level detection circuit 34 comprises a clamping circuit consisting of capacitors 226 and 227, and diodes 228 and 229. This clamping circuit receives a carrier signal and doubles the carrier voltage and clamps it below the floating ~round~ Capacitor 230 receives the output of the clamping circuit through resistor 233 and charges to an average DC level deter-mined b~ the received signal The charge on capacitor 230 holds the gate of FET 234 negative so that FET 234 is turned off when a carrier signal is present. When the carrier is interrupted, capacitor 230 rapidly discharyes through diode 232 and the voltage on the gate of FET 234 is raised so tha~ capacitor 238 is shunted to the output of amplifier 220. Thi action greatly reduces the charging time constant of capacitor 238 which tne quic~ly restores the baseline to its proper level.
In operation, bias voltages are produced from the carrier signal and applied to the active components of the respiration channel of the patient ~nit. Also, the carrier signal is dropped in voltage t'nrough steering diodes 204 and 205 to a level below .5 volts, and the current of the carrier signal is controlled to be approximately 600 ua by constant current diodes 202 and 203. The modulated respiration signal from patient 12 is detected by F~T 206 and AC
coupled by capacitor 214 to filter section 218 where the signal is smoothed. The filtered signal is amplified by operational amplifier 220, the output of which is sensed by capacitor 238 which maintains a voltage at the inverting input of amplifier 220 which is equal to the DC offset of the signal at the non-inverting amplifier input. The output of amplifier 220 is frequency modulated by voltage controlled oscillator 214 unless a patient lead is loose, in which case operational amplifier 240 disables the voltage controlled amplifier.
If the DC offset of the ECG data signal becomes too great, microprocessor 56 shuts off the carrier signal. The time during whicn the carrier signal is removed ;s not sufficient for capacitors 201 and 207 to discharge, the bias vol~ages are not removed from the active elements of the circuit. The capacitor 230 discharges thus turning on FET 234 thereby re~ucing the charging time constant of capacitor 238. After a predetermined time, the carrier is turned back on and circuit operation is resumed.
The ECG channel receives a signal on line 25 which is input to filter section 70. Clamping diodes 244 clamp ~he signal within predetermined limit.s of the floating ground and filte-r section 246 ~asses only the ECG signal and filters out the 100 Khz respiration signal and spurious signa:Ls due to, for example, defibrillation of the patient or the like.
The filtered ECG signal is received by amplifier 72 which comprises operational amplifier 248 and feedback resistor 250. Any DC offset in the ECG signal is compensated for by capacitor 252 which also receives the output of operational amplifier 248 through resistors 250 and 251. Finally, the o~tput of opera-tional amplifier 248 is passed to frequency modulation circuit 74 which comprises integrated circuit voltage control~ed oscillatvr 258 which can be an Intersil ICM
7555 integrated circuit which is appropriately biased.
Capacitor 252 can be connected directly to the output of operational amplifier 24~ by FET 254.
The gate of FET 254 is protected by back-to-back zener diodes 256 and is also connected to the output of level detection circuit 34.
In operation, the ECG signals are clamped by diodes 244 and filtered by filter section 246, The filtered ECG signals are amplified in operational amplifier 248 and the DC offset of the signal is reduced to zero by a charge on capacitor 252. When the offset on either the respiration channel or the 3CG channel becomes too great, microprocessor i6 causes ~ETs 234 and 254 to turn on thus reducing the charging time constants of capacitors 233 and 252.
Figure 3 shows the respira~ion channel of the signal analysis unit of the present invention.

The respiration channel is isolated from the patient unit by optical coupler 300 which receives the frequency modulated respiration signal on line 33 and optically couples this signal to transistor 302. The isolated signal from transistor 302 is passed to demodulation circuit 42 which comprises a low pass filter which includes capacitor 304, and an integrator which receives the output of the low pass filter. The integrator i5 formed from operational amplifier 306 and a feedback capacitor. The demodulated signal is connected tc the inverting input of amplifier 308 which is biased to provide gain to the signal and pass only the frequencies of interest.
A linear output from amplifier 308 is provided on line 311 to switch 45 which comprises a sinyle FET having an output connected to line 51. The output of amplifier 308 is also provided to logarithmic conver sion circuit 48. Circuit 48 includes operational amplifier 312, the output of which is fed back to its inverting input through steering diodes 314-317 and a ladder network comprising diodes 318 and resistors 320. Diodes 318 are connected in series and each has its anode connected to one terminal of a resistor 320.
The opposite terminals of resistors 320 are connected 2S together and to the anodes of diodes 316 and 317.
Temperature compensation is provided by thermistor 322. As can be seen, the output of operational amplifier 312 is fed back through the appropriate steering diodes and one or more of diodes 318 to the inverting input of the amplifier. Accordingly, the - ~5 -outPut o~ ampliier ~12 is t'ne logari~hm of the input.
As the input voltage increases, diodes 318 are incre-mentally included in the feedbac~ path to increase the range of the logarlthmic amplifier. The log output is ~rovided on line 51 through switch 46, which comprises a single FET, to output line 47.
Also connected to the output of transistor 302 is the loose lead detector 50 which comprises in~ut capacitor 324, clamping diode 326 and steering diode 328. Steering diode 328 is connected to the gate of FET 3320 A capacitor 330 is also connected to that gate. The output of FET 332 is presented to the microprocessor through line 55. A ~ener diode 334 provides input protection for the microprocessor.
Clearly, when no signal is present on line 33, the voltage to the gate of FET 332 increases and a low signal is passed along line 55 to the microprocessor.
The output of operational amplifier 306 is also passed to baseline detection circuit 52 which comprises comparators 336 and 338. The inverting input of comparator 336 is connected to a positive voltage source and the non-inverting input of comparator 338 is connect~d to a negative voltage source. The output of comparator 306 is connected to the non-invertin~ input of comparator 336 and the inverting input of comparator 338. Accordin~ly, if the level of the signal from comparator 306 goes above or below the level indicated by positive voltage source 340 or negative voltage source 342, respec-tively, a signal will be passed through dio~e 344 or 346 ! respectively, to transistor 348 to turn that transistor on and~pass a signal through line 53 to the microprocessor.

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Similarly, a second baseline detection circuit 5~ is connected to the output of operational amplifier 308. Baseline detector 54 is identical to offset baseline detector 52 and comprises comparators 350 and 352, voltage sources 354 and 356, and diodes 358 and 360. The output of circuit 54 sends a signal to the microprocessor through transistor 348.
The output of operational amplifier 308 is also passed to signal shaping circuit 43 which 0 comprises clamping diodes 362 and 363, capacitor 364, operational amplifier 366 and output transistor 370.
The signal to shaping circuit 49 is clamped by diodes 362 and 363, averaged by capacitor 364 and fed to amplifier 366. Amplifier 366 is connected with posi-tive ~eedback to produce a square wave with hysteresis each time that the respiration data signal goes above and below zero. The output of amplifier 366 is clamped to ~ 6 volts by diodes 36~ and 3~8, respectively. The output is also connected to the input of transistor 370 which passes the shaped signal to the micro-processor.
The ou~put of logarithmic amplifier 48 is fed back to the input of baseline correction circui~
58. Circuit 58 comprises resis~or 374 and an integra-tor comprising operational amplifier 372 and capacitor376. The output of operational amplifier 372 is connected to the non-inverting input of amplifier 308.
Accordingly, it can be seen that as the average signal from logarithmic amplifier 48 deviates from zero volts, the output of amplifier 372 will fluctuate and be added to the respiration data signal in amplifier 308 to return the baseline to zero volts.

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The charge on capacitor 376 can be ~aried more rapidly by turning on FET 398 of the fast correc-tion circu}t 63. FET 398 is effective to place a low value resistance 397 in parallel with resistor 374.
Accordingly, by turning on FET 398, the charge on the integrator of the baseline correction circuit 58 will be allowed to vary more rapidly. The gate of FET 398 is connected to the output of transistor 396, the base of which is connected through line 61 to an out~ut of the microprocessor. Accordingly, the micro processor can turn on FET 398 by an appropriate signal on line hl~ The output of transistor 396 is also passed to operational amplifier 399 which sends an output to turn off switches 45 and 46 by applying a negative signal to their gate inputs, and thus eliminate the linear respiration output on line 51 and the logarithmic output on line 47. Consequently, a display of the respiration signal will be stopped during the fast restore period.
Due to the extremely slow frequency of the respiration data signal, it may be useful to restore the baseline to the appropriate level by varying the charge on the integrator of baseline correction circuit 58 at a rate which is intermediate that provided by resistors 374 and 397. Accordingly, a medium slow charge circuit is included which comprises variable resistor cptical coupler 390 which has a variable resistance in series with resistor 392. This series resistance is in oarallel with resistor 374.
Acc~rdingly, when optical coupler 390 is turned on, a medium resistance is placed in parallel with high value resistance 374. Optical coupler 390 has a control input connected to ~he output of buffer 388.
The intput to buffer 388 is received from line 62 thro~gh transistor 378 and diode 384. Transistor 378 is also connected through resistor 386 ~o the input of an integrator comprising operational amplifier 380 and capacitors 382 and 383. Clearly, the rate of charge ca~ be varied by the microprocessor when a signal is ~rovided on line 62 to turn on transistor 378. The signal immediatelY causes a variation in the resistance of optical coupler 390, thus varying the time constant oE the baseline correction circuit.
Also, this time constant is reduced in accordance with the amount of time the signal is maintained by the integrator comprising amplifier 380 and capacit.ors 382 and 383. Back-to-back polar capacitors are used in the integrator to allow the in~egrator to output either polarity. The integrator charging time is rapid, so that the variable time constant rapidly changes toward fast under microprocessor control.
Integrator discharge time is slow so that the variable time constant slowly changes toward slow under the microprocessor control.
In operation, when a signal is received by optical coupler 300, that signal is electrically isolated from the patient unit and passed to the demodulation circuit which includes capacitor 300 and integrator 306. If no signal is receiYed, the microprocessor is notified through FET 332 and line 55, and a loose lead alarm is actuated.
The of f set level of the demodulated output is compared in comparators 336 and 333 to maximum and minimum permissible levels and, if the baseline is not within the prescribed range, transistor 348 is actuated to notify the microprocessor.

.. . .

The demodulated respiration data siynal is amplified in amplifier 308. The offset of tl~e amplifier output is again checked with permitted maximum and minimum values and, if outside the acceptable range, the microprocessor is notified by a signal from transistor 348. The ECG dat~ signal is also directly passed to positive feedback amplifier 366 which acts as a zero crossing detector to notify the micropro~essor of each respiration e~ent through transistor 370. Further, the amplified ECG data signal is passed to switch 45 which, if fast restoration of the baseline is not being performed, passes the signal to linear respiration output lead 51. The linear output is also passed to logarithmic ampli~ier 4 a which logarithmically amplifies the signal and passes it through switch 46 to output line 47. If a display of the signal is required, an appropriate display device can be attached to either line 47 or line 51.
If fast baseline restoration is required, a signal is produced on line 61 which turns on FET 398 to place resistor 397 in parallel with resistor 374 thus reducing the time constant of the baseline correction circuit. After fast baseline restoration ~5 is complete, medium slow variable baseline restoration i~ effected by a signal on line 62 which controls the resistance of optical coupler 390 which is placed in series with resistor 392 and which combination is placed in parallel with resistor 374. The medium slow variable baseline rate is controlled by the output of the integrator circuit comprising operational amplifier 380 and capacitors 382 and 383.

., Figure 4 shows the ECG channel of the signal analysis unit. An isolation device 76 comprising optical coupler 400 receives frequency modulated ECG
signals on input line 73 and passes an output through transistor 402 to the ECG channel. The ECG channel includes a frequency demodulator comprising a low pass filter which includes capacitor 404, and an integrator comprising operational amplifier 406 with a capacitor in its ~eedback loop. The demodulated signal is ~assed to an amplifier comprising operational amplifier 408 having a feedback network connected to its inverting input so that the amplifier will pass only frequencies associated with the ECG signal. I'he output of operational amplifier 408 is passed to switch 82 comprising a single FET 410 which, when on, passes the sisnal to output line 100. The output of amplifler 408 is also passed to automatic gain control circuit 84. Automatic gain control circuit 84 includes operational amplifier 411 which is biased as a linear amplifier and also includes the variable resistance p~rtion of a variable resistance optical coupler 412 in i~s feedback loop. The output of operational amplifier 411 is AC coupled through capa~itor 414 to the input of transistor 418. The input of transistor 2~ 418 is clamped to ground by diode 416. Transistor 418 feeds an integrator comprising operational amplifier 420 and capacitor 4220 The charge on capacitor 422 is clamped in a forward direction by diode 424.
Capacitor 414 and diode 416 act to clamp the signal ~0 input to transistor 14 to ground. Accordingly, the integrator reacts to the peak amplitude of the ECG
signal and, when the amplitude increases, the output .

of integrator 420, 422 alsc increases thus decreasing the resistance of optical coupler 412 and thereby decreasing the gain of amplifier 411.
The output of amplifier 411 has a constant s peak amplitude and is passed through switch 101 to output line 102 anæ to signal shaping circuit 86.
Circuit 85 comprises a filter portion 460 which passes the R wave portion of the ECG data signal. The R wave portion is received by the non-inverting input of o~erational amplifier 462 and compared with a voltage source 464. If the R wave is greater than the vol~age set by source 464, a square wave output is passed through transistor 464 to line 87 and to the micro-processor. A linear ECG signal is made available at the output of FET 410 on line 100, and a gain controlled signal is made available on line 102.
The output o~ operational amplifier 408 is fed back through line 91 to baseline correction circuit 90. Baseline correction circuit 90 comprises a high value resistance 428 which feeds ~he input of an integrator comprising operational amplifier 430 and capacitor 432. The output of integrator 430, 432 is fed to the non inverting input of operational amplifier 408. Accordingly, as the DC offse~ of the output of amplifier 408 changes, the charge on integrator 430, 432 changes also and ~ullifies the DC
offset. Fast baseline correction circuit 93 comprises ~ET 434 which, when turned on, connects lo~ value resistance 435 in parallel with resistance 428. FET
434 is turned on by transistor 436 in response to a signal from the microprocessor on line 620 Line 62 is clamped by diodes 438 and 439. Line 91 is also connected to an of~set detection circuit 88 which ,,. ~ :

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comprises comparators 442 and 444 which have their non-inverting and inverting inputs connected respec-tively to line 91. Capacitor 440 is connected to line 91 also for averaging the signal applied to co~
parators 442 and 444. Comparator 442 compares the signal on line 91 to a positive voltage provided by source 446, and comparator 444 compares the signal on line 91 to a nesative voltage provided by voltage source 448. Comparators 442 and 444 provide outputs if the offset of the signal on line 91 is above or ~elow the voltages set by sources 446 and 448, respec-tively. These outputs are passed through diodes 450 and 452, respectively, to transistor 454 which sends a si~nal through line 89 ~o the microprocessor.
lS In operation, optical coupler 400 feeds transistor 402 which provides an isolated frequency modulated ECG signal to demodulator 78. The signal is low ~ass filtered and integrated in the demodulator and passed to operational amplifier 40~ in which the ~o signal is amplified. Any offset in the signal is fed back through line 91 to charge integrator 430, 432 which then reduces tbe offset at amplifier 408. If fast baseline correction is required, a signal is provided at line 62 which turns on FET 434 and reduces 25 the time constant of the integrator. At the same time, a signal is provided on line 94 to switch off switches 82 and 101. The signal on line 91 is also sensed by amplifiers 442 and 444 which feed a signal through line 89 if a DC offset is too great. In response to such a signal, the microprocessor may begin fast offset correction.

~z~

When FET 410 of switch 82 is ~urned on, the baseline corrected signal is oassed to line 480 which may be connected to a dis~lay device or the like. The signal is also passed to the automatic gain control circuit. Initially, the signal is received by amplifier 411 which is controlled to have a maximum gain. This high gain signal is passed through tran-sistor 418 to integrator 420, 422 which reduces ~he gain of amplifier 411 to the desired amount. The output of integrator 420, 422 then con~rols the gain of amplifier 411 so that the peak amplitude o~ the ECG
signal is maintained at a desired level. The gain controlled signal is filtered at filter 460 and squared in differential circuit 462 and passed through line 87 to the microprocessor~ If the signal at the microprocessor disappears~ a signal can be passed through line 65 and diode 426 to rapidly decrease the output of integrator 420, 422 thus increasing the gain of amplifier 411 to its maximum in an attempt to restore the siqnal level.
Figure 5 shows the carrier genera~or 18 of the present invention. 5enerator 18 comprises an astable multivibrator formed from transistors S00, S0~ capacitors 504, 508, and resistors 506, 510. The collectors of transistors 500, 502 are fed from a S volt supply through opposite halves of the center tap primary of transformer 20. The astable multivibrator operates in a conventional manner. When the carrier si~nal is to be removed, the micro~rocessor sends a ground signal through line ~0 to transistor 514.
Tr~nsistor 514 is turned on, thus turning on FET 512.
FET 512 shorts out the bases of transistors ;00, 502 to stop the operation of th- mul~ivibrator.

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~ igures 6a, 6b, 7a, 7b and 7c show flow charts for the main and interrupt service programs used by microprocessor 56 in the present invention.
The interrupt service program is executed once every millisecond. This program is entered at step 700, at which ti~e certain initialization functions are performed as would be obvious ~o one of ordinary skill in the art. At step 702, registers ECGNEW and RESNEW are incremented. At step 660, the setting of apnea switch 96 is stored in a register called SW. At step 662, the contents of SW are compared with the con~en~s of a register called APNEA
which stores the last valid setting of the apnea switch. If the apnea switch setting has been chanqed, step 664 passes control to step 666 which determines whether the mismatch FLAG has been set. If the mismatch FLAG has been set, control passes to step 668 which se~s a tamper alarm. If the mismatch FLAG has not been set, control passes to step 670 which deter~
mines whether a ground signal is present on line 99 indicating that reset switch 97 has been actuated. If the reset switch is depressed, control passes to step 672 which stores the new apnea switch setting in ~PNE~. If the reset switch is not depressed, control passes to step 674 which sets ~he mismatch FLAG.
If, at step o64, ~he apnea switch has not been changed, control passes to step 676 which clears the mismatch FLAG. Control is then passed to step 578 which dtermines whether reset switch 97 is depressed.
If switch 97 is depressed, control passes to step 680 which clears the ~amper alarm.

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As can be seen from the above desciption, this portion of the program causes an alarm to be set if the apnea switch setting has been changed without the reset switch being depressed and the changed setting is held for two cycles of the interrupt service routine. On the first cycle, the mismatch FLAG
is set at step 674 and on the second cycle, the tamper alarm is set at step 668. However, if the reset switch is depressed when the apnea switch setting is changed, steps 670 and 672 cause the new setting to be stored as a valid apnea time. Once a tamper alarm has been set, in order to clear this alarm, the apnea switch 96 must be returned to its original, valid setting and reset switch 97 ~lust be depressed. This lS action causes control to ~e passed from step 664 through steps 676 and 678 ~o step 680 which clears the tamper alarm.
After steps 668, 672, 674 or ~80, control is passed to step 704. At step 704, a clock register is checked to determine whether an increment of 128 mi~liseconds has occurred. For each increment of 128 milliseconds, control passes to step 706, at which time lines 53 and 89 are checked t~ determine whether the respiration or ECG baseline signals are excessive.
If the signals are off scale, a register named OFF is set to be equal to 255 in step 708. Thereafter, the contents of OFF are decremented by 1 in step 710. If the ECG or respiration siynals are not off scale, OFF
i9 also decremented by l in step 710. Accordingly, it will be seen that once OFF is set to 255, it takes approxima~ely 32 seconds for the contents of that :

32~

register to be decremented to zero, since step 706 is performed approximately once every eighth of a second.
In step 712, the contents of OFF are checked. If the CQntents are greater ~han 253, control passes to step 714, which controls the carrier kil~ circuit. The carrier is killed through line 60. If the contents of OFF are less than 253, control passes to step 716, which restores the carrier. Control is then passed to step 718 where the contents of OFF are checked to see if they are ~reater than or equal to 240. If the contents of this register are greater than or equal to 24b, step 7~0 sets the fast ECG baseline by passing a signal through line 61 Otherwise, the fast E~G base-line is cleared at step 722. Con~rol is then passed to step 72~ which checks to see if the contents of OFF
are greater than or equal to 232. If greater, control is passed to step 726, which sets the fast respiration baseline by passing a signal through line 61.
Otherwise, the fast respiration baseline is cleared at step 728. Control is then passed to step 730 where OFF register is checked to see if its contents are greater than zero. If greater, the medium respira~ion baseline is set at step 732 by passing a signal through line 620 Otherwise, the medium respiration baseline is cleared in step 734. Control passes from ste~ 732 or 734 to step 736. Likewise, if a 128 millisecond increment has not been determined at step 704, control passes directly to step 736. At s~ep 736, input line 87 is checked to determine if a square wave input is present, indicating a new ECG event. If an event has occurred, control passes to step 738 where the contents of the ECGNEW register pair is 2~

~oved to the ECGOLD register pair, thus providing a determination of the increment between the last two ECG events in the ECGOLD register pair. The ECG event flag is set and the ECGNEW register pair is cleared.
If no new ECG event has occurred or ~fter the completion of step 738, control passes to step 740.
Step 740 checks line 57 to determine whether a new respiration event has occurred in the form of a square wave signal on line 57. If a new respiration event has occurred, control passes to step 742 where the contents of the RESNEW register pair is moved to the RESOLD register pair, thus providing an indication of the interval between the last two respirations. The RES event flag is set and the RESNEW register pair is cleared. After completion of step 742, or if no new respiration event has occurred, control returns to the main program.
Figures 6a and 6b set forth the main program. The main program is initialized in step 600 and control is then passed to step 602 where the ECG
event flag is checked. If the ECG event flag is set, control passes to step 604 where the ECG flag is set and the ECG event flag is reset. If the ECG event flag is not set, control passes to step 606 where the ECG flag i5 cleared. Control is then pased to step 608 where the respiration event flag is checked. If the respiration event flag is set, control passes to step ol0 where the respiration flag is set and the respiration event flag is resetO If the respirakion event flag is not set at step ~08, control is passed to step 612 where the respiration flag is cleared.

, Control is then passed t~ step 614 where the ECG flag is checked. If the ECG flag is set, con~rol passes to step 616 which calculates the ~CG rate by di~iding the contents of the ECGOLD register pair by 60,000 to provide a beats-per minute indication. Step 618 determines whether the beats per minute is greater than 255. If the beats per minute is less than 255, the ECG rate register is provided with the actual ECG
rate. Otherwise, the ECG rate register is set at 255 in step 620. In step 622, the new ECG rate is moved into a bucket brigade which stores the last 8 ECG
rates. At step 624, the average ECG rate is deter-mined by averaging the contents of the 8 buckets of the bucket brigade. At step 626, the most recent ECG
rate is tested, This is accompli.shed by dividing the most recent ECG rate by a constant to provide an interval number and dividing the average ECG rate by a constant to provide an interval number. In step 628, the recent ECG interval is divided by the average ECG
interval. The result is compared with 2.25.
Returning for a moment to step 614, if the ECG flag is not set, the test in step 626 is performed on the previous most r~ent and average ECG rates. In step 628, if the recent ECG inter~al is equal to or greater 2S than 2~25 times the average ECG interval, control passes to step 630, where ~he fast ~GC is set by passing a slgnal on line 65. Otherwise, t~e fast AGC
is cleared at step 632. Control is then passed to step 634 which checks to see if the respiration flag has been set. If the respiration flag has been set, control passes ~o step 636, which calculates the respiration rate by dividing the respiration interval in the RESOLD register pair ~y 60,000. If it is ::`

determined in step 638 that the respiration rate is greater than 250, the res~iration rate register is set at 250 in step 640. Otherwise, the actual respiration rate is set in the respiration rate register. Control is then passed to step 639 where the new respiration rate i5 added to a ~ucket brigade. In step 641, the respiration average rate is determined. Control i~s then passed to step 642 where the average respiration interval is compared to the contents of APNEA divided ~y 2. As will be recalled, APNEA contains the last valid setting of APNEA switch 96. This wili be a value of approximately 5 or 7 seconds. Of course, the values determined by AP~EA switch 96 can be set at any amount desired. If the respiration flag has not been set at step 634, control passes directly to step 642.
If it is determined that the average respLration interval at step 642 is greater than APNEA divided by ~, control passes to step 644, where OFF is incre-mented by 2. Control then passes to step 646, where the loose lead input is checked on line 55. If a loose lead condition is present, control is passed to step 648, which sets an alarm register. Otherwise, control is passed to step 653, which resets the alarm register. Control is then passed to step 652, which returns control to the start block 654, ~hich then passes control to step 602 again.
As will be understood from the foregoing explanation, every millisecond, the interrupt service routine of Figures 7a, 7b and 7c tests for a new ECG
event and a new respiration event in step 736, and 740, respectively When a new ECG event occurs, a number is stored in the ECGOLD register pair indicating the time interval in milliseconds between the last two ECG events. In like manner, a number is stored in the RESOLD register pair each time a new res~iration event occurs. This number indicates in milliseconds ~he time interval between the last two respiration events. Upon each multiple of 128 milliseconds, input lines 53 and 89 are checked to determine if an off scale indication from baseline detection circuits 52, 54 or 88 is present. If an off scale indication is present, the carrier generator 18 is turned off through line 60 for a minirnum of approximately one-eighth of a second at step 712 and 714. During this time, rapid baseline restoration is effected by circuits 36 and 75 in the patient unit.
Also, steps 718 and 720 cause the ~ast baseline circuit ~3 to be operative for approximately 2 seconds, duriny which time rapid baseline correction of the ECG data signal is performed by circuit 90. Also at the same ~ime, steps 7~4 and 726 cause a rapid restoration of the respiration baseline through circuit 63 for approximately 3 seconds. Finally, medium-slow circuit 64 is set for approximately 32 seconds so that, after the rapid baseline correction of the respiration data signal is completed, medium-slow baseline correction continues for approximately another 30 seconds. The program then tests for new ECG and new respiration events as discussed above. Control then returns to the main program of Figures 6a and 6b.
As can be seen from the discussion above, the main program is operative to first determine if a new ECG event or new respiration event has been sensed , .

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~y the interrupt program. If no new event has occurred, the program, at steps 626 and 628, tests to determine whether the last ECG interval is greater than the average interval by a predel:ermined amount.
s If the interval is greater, the fast AGC in2ut line 65 would have been activated to increase the gain of AGC
circuit 84 at step 630. If no new EC`G event has occurred, the fast AGC input is held activated. The program then tests at step 642 to determine whether the previo~s respiration interval is greater than a normal respiration. If the last interval was greater, the medium~slow baseline respiration circuit 64 would have been activated by incrementing the OFF register by 2. Accordingly, if no new respiration event has occurred, ~he OFF register is again incremented by 2.
The proyram then checks to determine whether a loose lead has occurrred. If so, an alarm is sounded. If not, the prQgram returns to step 654.
If an ECG event has occurred, the program calculates a new ECG average for the last 8 occurrences in step 624 and tests the most recent ECG against the average in steps 626 and 628. Again, if the most recent ECG interval is greater than the averge by a predetermined amoun~, the fast AGC circuit is activated at step 630.
In like manner, if a new respiration event has occurred, a new respiration average is determined at step 639 and the average respiration interval is tested at step 642. If the interval is greater than set by APNEA switch 96 divided by 2, the OFF register is incremented by 2 so that the medium-slow respira-tion baseline correction circuit 64 is activated.

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Figure 8 shows a po~er supply to be used with the present invention. The power supply 95 includes a standard three-pronged plug 800 which can be connected to a wall socket. Lines 802 and 804 receive current from plug 800 and pass the current through two poles of a three-pole switch 806 ~nd a transformer 808. The current is rectified in diodes 810 and 812. The rectified current is passed through line 814 and the third contact of switch 806. The rectified current is riltered at capacitor 816.
Alternatively, an external battery can be connected to the power supply through terminals 818 and diode 820.
The filtered current from capaci~or 816 is passed to unregulated supply line 822. A first integrated circuit voltage regulator 824 provides regulated positive S volts on line 826 for use in ~he analog circuits of the invention. A second integrated circuit voltage regulator 828 provides regulated +5 volts on line 830 for the digital circuits of the present invention. A negative 5 volt supply is provid~d by voltage mirror circuit 832 which produces an outp~t on line 834 which i5 the negative of an input received on line 836. The input on line 836 is controlled by transistor 838 which receives an input from line 82~. The conduction of transistor 838 is controlled by operational amplifier 840 which is biased to provide a gain of approximately 100. The non-inverting input of amplifier 840 is connected to ground. The inverting input of that amplifier has one lead connected through resistor 842 to the positive 5 volt supply on line 826. A second lead is connected ~ j ~v~

through resistor 844 to the negative output of voltage mirror circuit 832~ Accordingly, the voltage at the inverting input of amplifier 840 is the sum of the voltages on lines 826 and 834. The output of amplifier 840 changes in accordance with the sum of the voltages on lines 826 and 834, and is connected through line 846 to the base of translstor 838.
Consequently, the conduction of translstor 838 is increased or decreased in accordance with the output of circuit 832 compared with the voltage on line 826.
Accordingly, the negative 5 volt source is regulated in this manner.
A start~up transistor 848 has its collector connected to the base of transistor 838. ~ransistor 848 is momentari].y turned on by capacitor 850 when switch 806 is closed. This initi~lly turns on transistor 838 to ensure that an initial minus voltage is provided without delay.
The output of amplifier 840 is also passed through line 841 to the base of transistor 843. If the output of circuit 832 is correct, the output of amplifier 840 is slightly positive thus turning on transistor 843 through line 841. The collector of transistor 843 is connected to the base of transistor 845 through line 847. The collector of transistor 845 is connected to line 849 which actuates an alarm (not shown) when line 849 goes low. Clearly, when tran sistor 843 conducts, line 847 goes low and transistor 845 is non-conducting thus holding line 849 high. In the event that either too much current is drawn from circuit 832 through line 834 or the voltage on line 822 drops, the voltage on line 834 changes thus : ,~

incre2sing the input to amplifier 840. The o~tput of amplifier 840 is accordingly reduced to increase the conduction of transistor 838. Accordingly, the voltage on line 841 is similarly reduced. When the voltage on line 841 falls below approximately .7 volts, transistor 843 ceases to conduct and the voltage on line 847 increases. This turns on tran-sistor 845 which reduces the voltage on line 849.
Consequently, a low voltage alarm connected to line 849 will be actuated.
Supply voltage circuit 95 also contains a tampering alarm which is activated in the event that switch 806 is opened by mistake. The tampering alarm comprises flip flop 852 which has its non-inverted output connected to transistor 85~. When flip flop 852 is set, transistor 854 is turned on thus sending a ground signal through line 855 to an appropriate alarrn. The set input o~ flip flop 852 is connected through line 856 and line 857 to capacitor 858.
Capacitor 858 is connected to the collector of transistor 859. Also, line 857 is connected to ground through filter capacitor 860, resistor 861 and clamping diode 862. The collector of txansistor 859 is also connected through resistor 863 to line 864.
Line 864 is supplied with positive voltage from unregulated supply line 822 through diode 865 or from a 9 volt battery connected to terminal 866 throush diode 867. Line 864 also provides positive bias voltage for flip flop 852. The base of transistor 859 is connected through a voltage divider to line 822.

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Line 857 is also connected to the cat:hode of diode 870, the anode of which is connected to the base of transistor 871. The base and collect:or of transistor 871 are connected through bias resist:ors to line 873 which is connected to unregulated suE~ply line 822.
The collector of transistor 871 is also connected through line 874 to the reset input of flip flop 852.
In operation, as long as unregulated supply voltage is provided on line 822, transistor 859 is held in a conducting state and bias voltage is supplied to flip flop 852 through line 864 and diode 865. If the unregulated supply voltage is removed, as when switch 806 is opened, diode 865 becomes non-conducting and diode 867 conducts voltage to line 864.
The base voltage of transistor 8S9 is removed and the transistor becomes non-conducting. The battery voLtage from terminal 866 ~hrough diode 867, line 864 and resistor 863 causes an increase in voltage at ~he collector of transistor 859. This voltage increase is differentiated by capacitor 858 which causes a posi-tive spike to pass through line 856 to the set input of flip flop 852. Flip flop 852 is set and thus passes a positive signal rom its non-inverted output to transistor 854. Transistor 854 becomes conducting and line 855 goes to ground. This signal may be sent to the microprocessor of the sys~em or be used to actuate any appropriate alarm.
Line 856 and the set input of flip flop 852 are also connected to reset switch 97 through line 98.
3G Cleaxly, depressing reset switch 97 after power has been interruped has no effect~ Such depression of .

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switch 97 merely provides a ground signal to the set input of fli~ flop 852 and, through diode 870 to the base of transistor 871. However, since the bias voltage on line 873 has been removed, no signal is sent to the reset input of flip flop 852 through line 874. In order to reset flip flop 852, the supply voltage must be reestablished, as through closure of switch 806. Once bias voltage has been returned to transistor 871, if switch 97 is depressed, a grouncl signal through diode 870 to the base of transistor 871 causes that transistor to become non-conducting.
Accordingly, line 874 goes high sending a positive signal to the reset input of flip flop 852 resetting the flip flop.
Clearly, if i~ is desired to turn the power off, reset switch 97 should be depressed at the same time as switch 806 is opened. This causes the set pulse on line 856 to be shunted directly to ground and flip flop 852 never is set.
Figure 9 shows a circuit which can be used to test the monitor of the present invention. The test circuit produces a 100 millisecond duration, .2 millivolt amplitude pulse at the rate o 154 per minute to sitnulate ECG. It also generates a .2 ohm a~plitude, one second duration pulse at the rate of 77 per minute to simulate thoracic impedance ~ariation due to respiration.
The circuit comprises an integrated circuit oscillator 900 which ~an be an In~ersil ICM 7555 oscillator biased to yenerate pulses at a rnultiple of 154 per minute. The output of oscillator 900 is passed through line 902 to counter circuit 904.
Counter 904 can be a Motorola Mudel MC14040 counter.

8;~

Two outputs fro~ counter 904 are provided on lines 906 and 908. The output on line 906 is at a frequency twice that of the output on line 908. The frequency of the output on line 906 should be 154 pulses per minute, and the output on line 908 should be 77 pulses per minute. Line 906 is connected to capacitor 910 which differentiates the pulses and provides spikes to tne base of transistor 912 through a vol~age divider network. The output of transistor 912 is taken from its collector through resistor 914 which may be a 200K ohm resistor. Accordingly, it can be seen that the output of transistor 912 comprises a olurality of spikes which are similar to the R wave por~ion of an ECG wave. Thes~ spikes are passed through resistor 914 to line 916 at a rate of 154 per minute.
Line 908 is connected through resistor 918 to the gate of FET 920. A 499 ohm resistor 321 is connected between the drain of FET 920 and line 916.
A 10 ohm resistor 922 is connected from the source of FET 920 to line 916. A 100 ohm resistor 924 is connected between the source of FET 920 and output terminal 930. Accordingly, it will be seen that FET
920 will be turned on at a rate of 77 times pe~ minute in accordahce with the pulses received from line 908.
Each time FET 920 is turned on, resistor 921 is placed in parallel with resistor 922 thereby bringing the combined resistance to a total slightly less than 10 ohms. This resistance change is approximately equal to .2 ohms~ Consequently, it can be seen that the resistance variation between lines 916 and 930 will be approximately .2 ohms at a rate of 77 times per minute.

It will be understood that lines 930 and 916 will be connected to the inp~t of circuit 14 in place of the patient connected leads and ground lead 932 should be connected to the ground of circuit 14. In this manner, simulated respiration and ECG waves will be processed by the monitor of the present invention.
Out~uts from the monitor 10 can be observed to determine whether the system is operating properly.
The foregoing description is set forth for the purpose of illustrating the invention but is not meant to limit the scope thereof in any way. Clearly numerous modifications, additions, and other changes can be made to the presen~ invention without departing from the scope thereo~ as set forth in the appended claims.

Claims (53)

THE EMBODIMENTS OF THE INVENTION IN WHICH AN EXCLUSIVE PROPERTY
OR PRIVILEGE IS CLAIMED ARE DEFINED AS FOLLOWS:

1. A patient monitor comprising:
a patient unit comprising:
a probe connected to receive a carrier signal, said probe being adapted for connection to the body of a patient to be monitored, whereby said carrier signal is modulated in accordance with respirations of said patient to produce a modulated carrier signal;
carrier detection means connected to receive said modulated carrier signal and produce a demodulated respiration signal;
amplifier means for amplifying said demodulated respiration signal to produce an amplified respiration signal;
frequency modulation means for receiving and frequency modulating said amplified respiration signal to produce an FM respiration signal; and an analysis unit comprising carrier generation means for producing said carrier signal;
a frequency demodulation circuit for receiving said FM respiration signal and demodulating said FM respiration signal thereby producing a respiration data signal; and an output circuit connected to receive said respiration data signal, amplify and level shift said respiration data signal, and output said respira-tion data signal.

2. The monitor as set forth in claim 1 and further including isolation means for electrically isolating said patient unit from said analysis unit.

3. The monitor as set forth in claim 1, wherein said patient unit includes a baseline correction circuit means for sensing the DC level of said demodulated respiration signal and adding or subtracting a DC signal to said demodulated respira-tion signal in response to said sensed DC level.

4. The monitor as set forth in claim 3, wherein said baseline correction circuit means includes a capacitor which is charged in response to said sensed DC level.

5. The monitor as set forth in claim 4 and further including means for discharging said capacitor when the DC level of said respiration data signal reaches a predetermined amount.

6. The monitor as set forth in claim 5, wherein said discharging means comprises means in said analysis unit for deactivating said carrier generation means to stop the production of said carrier signal, and means in said patient unit for sensing the lack of carrier signal and discharging said capacitor in response thereto.

7. The monitor as set forth in claim 1 and further including a baseline correction circuit means in said analysis unit for sensing the DC level of said respiration data signal and adding or subtracting a DC
signal to said respiration data signal in response to said sensed DC level.

8. The monitor as set forth in claim 7, wherein said baseline correction circuit means contains a capacitor which is charged in accordance with said sensed DC level.

9. The monitor as set forth in claim 8 and further including means for varying the rate of charging said capacitor.

10. The monitor as set forth in claim 9, wherein said varying means includes means for rapidly discharging said capacitor when the DC level of said respiration data signal is above or below predeter-mined limits.

11. The monitor as set forth in claim 10, wherein said rapid discharge means includes a programmed microprocessor.

12. The monitor as set forth in claim 10, wherein said rapid discharge means includes a pair of differential amplifier circuits connected to receive said respiration data signal and produce output signals when said respiration data signal is above or below upper and lower limits, respectively.

13. The monitor as set forth in claim 9, wherein said varying means includes means for slowing charging and discharging said capacitor.

14. The monitor as set forth in claim 9 and further wherein said analysis unit includes a switch for connecting said frequency demodulation circuit to said output circuit when closed and disconnecting said frequency demodulation circuit from said output circuit when open, and wherein said rapidly varying means is operative for opening said switch when said DC level of said respiration data signal is above or below said predetermined limits.

15. The monitor as set forth in claim 1 and further wherein said patient unit includes a loose lead detector means for sensing the presence of said demodulated respiration signal and deactivating said frequency modulation means when no demodulated signal is present.

16. The monitor as set forth in claim 15 and further wherein said analysis unit includes a loose lead detector for sensing the presence of said respiration data signal and producing an output signal when no respiration data signal is present.

17. The monitor as set forth in claim 1 and further wherein said carrier detection means comprises a synchronous detector.

18. The monitor as set forth in claim 17, wherein said carrier detection means further includes a low pass filter connected to the output of said synchronous detector to pass frequencies below said carrier signal frequency.

19. The monitor as set forth in claim 1 and further including an ECG sensing circuit connected to said probe for sensing ECG signals produced by said patient.

20. The monitor as set forth in claim 19 wherein said ECG sensing circuit includes a filter contained in said patient unit for passing frequencies associated with an ECG signal.

21. The monitor as set forth in claim 20 and further including ECG amplifier means for amplifying said frequencies passed by said filter to produce an amplified ECG signal, and ECG frequency modulation means connected to frequency modulate said amplified ECG signal to produce an FM ECG signal, said ECG
amplifier means and said ECG frequency modulation means being contained in said patient unit, and an ECG
frequency demodulation circuit contained in said analysis unit for demodulating said FM ECG signal to produce an ECG data signal, and an ECG output circuit contained in said analysis unit for receiving said ECG
data signal, amplify, level shift and output said ECG
data signal.

22. The monitor as set forth in claim 21, wherein said ECG output circuit further includes an automatic gain control circuit for controlling the amplitude of said ECG data signal to be within predetermined limits.

23. The monitor as set forth in claim 22, wherein said automatic gain control circuit includes a gain controllable amplifier, and a gain control circuit connected to receive the output of said gain controllable amplifier and increase or decrease the gain thereof in response to said received output thereof.

24. The monitor as set forth in claim 23, wherein said gain controllable amplifier includes a variable resistance optical coupler contained in a feedback path, and said gain control circuit being connected to increase or decrease the resistance of said optical coupler in response to the received output of said gain controllable amplifier.

25. The monitor as set forth in claim 24, wherein said gain control circuit includes an integrator circuit having an input connected to receive said output from said gain controllable amplifier, and an output connected to said optical coupler.

26. The monitor as set forth in claim 21, wherein said patient unit is electrically isolated from said analysis unit.

27. The monitor as set forth in claim 21 and further including an ECG baseline correction circuit means for sensing the DC level of said amplified ECG signal and adding or subtracting a DC
signal to said amplified ECG signal in response to said sensed DC level.

28. The monitor as set forth in claim 27, wherein said ECG baseline correction circuit means includes a capacitor which is charged in accordance with said sensed DC level.

29. The monitor as set forth in claim 28 and further including a respiration baseline correction circuit means for sensing the DC level of said amplified respiration signal and adding or subtracting a DC signal to said amplified respiration signal in response to said sensed DC level.

30. The monitor as set forth in claim 29, wherein said respiration baseline correction circuit means includes a capacitor which is charged in accordance with said sensed DC level.

31. The monitor as set forth in claim 30 and further including means for discharging said capaci-tors in said ECG and respiration baseline correction circuit means in response to the DC level of said respiration data signal reaching a predetermined amount, and in response to the DC level of said ECG
data signal reaching a predetermined amount.

32. The monitor as set forth in claim 31, wherein said discharging means comprises means in said analysis unit for deactivating said carrier generation means to stop the production of said carrier signal, and means in said patient unit for sensing the lack of carrier signal and discharging said capacitors in response thereto.

33. A patient monitor comprising:
carrier generation means for producing a carrier signal;
a probe connected to receive said carrier signal, said probe being adapted for connection to a patient to be monitored, whereby said carrier signal is modulated in accordance with respirations of said patient to produce a modulated carrier signal;
carrier detection means connected to receive said modulated carrier signal and produce a demodulated respiration signal;
amplifier means for amplifying said demodulated respiration signal and outputting an amplified respiration signal;
an ECG filter connected to receive an ECG signal produced by said patient and filter said ECG signal from said carrier signal;
ECG amplifying means for receiving said ECG signal, and amplifying and outputting said ECG
signal;
respiration baseline correction circuit means for sensing the DC level of said respiration signal and adding or subtracting a DC signal to said respiration signal in response to said sensed DC
level;
ECG baseline correction circuit means for sensing the DC level of said ECG signal and adding or subtracting a DC signal to said ECG signal in response to the sensed DC level thereof; and a baseline reset circuit including level sensing means for sensing the DC level of said respiration signal and the DC level of said ECG signal and deactuating said carrier generation means when either of said DC levels becomes excessive, and control means for sensing the deactuation of said carrier generation means and controlling both the ECG baseline correction circuit means and the respiration baseline correction circuit means to return to an original condition.

34. A patient monitor according to claim 19 and further including a logarithmic compression circuit having an input connected to receive said ECG signal and an output for providing a signal which is proportional to the logarithm of said received ECG signal in order to reduce the dynamic range of said ECG signal.

35. The patient monitor as set forth in claim 34, wherein said logarithmic compression circuit includes an operational amplifier and a plurality of diodes connected in a feedback network on said operational amplifier.

36. The patient monitor as set forth in claim 1 and further including an actuatable reset switch and a power supply which can be turned on or off, and alarm circuit means for producing an alarm signal when said power supply is turned off if said reset switch is not actuated.

37. The patient monitor as set forth in claim 36, wherein said alarm circuit means further includes inhibit circuit means for eliminating said alarm signal only when said power supply is turned on and said reset switch is actuated.

38. The patient monitor as set forth in claim 36, wherein said power supply further includes low voltage alarm circuit means for producing an alarm signal when the output of said power supply falls below a predetermined limit.

39. The patient monitor as set forth in claim 36, wherein said alarm circuit means includes means for inhibiting said alarm signal if said reset switch is actuated when said power supply is turned off.

40. The patient monitor as set forth in claim 1 and further in combination with a test signal generation means for producing a simulated respiration wave form, said test signal generation means having output terminals for connection to said patient unit.

41. The combination as set forth in claim 40, wherein said test signal generation means comprises an oscillator, a fixed resistance, and a semiconductor component having an input connected to an output of said oscillator and having output terminals connected across said fixed resitance.

42. The patient monitor as set forth in Claim 19 and further including a test signal generation circuit means for producing simulated ECG and respiration wave forms, said test signal generation means having outputs for connection to said patient monitor.

43. The combination as set forth in claim 42, wherein said test signal generation means comprises an oscillator, a counter connected to said oscillator, said counter having a first output providing pulses at a first frequency and a second output for providing pulses at a second frequency, means for differentiating signals received from one of said outputs, and semiconductor means actuated by said differentiated signals, and variable resistance means connected and responsive thereto to produce a variation in resistance.

44. A patient monitor according to claim 33 and further including a logarithmic compression circuit having an input connected to receive said ECG signal and an output for providing a signal which is proportional to the logarithm of said received ECG signal in order to reduce the dynamic range of said ECG signal.

45. The patient monitor as set forth in claim 44, wherein said logarithmic compression circuit includes an operational amplifier and a plurality of diodes connected in a feedback network on said operational amplifier.

46. The patient monitor as set forth in Claim 33 and further including a test signal generation circuit means for producing simulated ECG and respiration wave forms, said test signal generation means having outputs for connection to said patient monitor.

47. The combination as set forth in claim 46, wherein said test signal generation means comprises an oscillator, a counter connected to said oscillator, said counter having a first output providing pulses at a first frequency and a second output for providing pulses at a second frequency, means for differentiating signals received from one of said outputs, and semiconductor means actuated by said differentiated signals, and variable resistance means connected and responsive thereto to produce a variation in resistance.

48. The patient monitor as set forth in claim 33 and further including an actuatable reset switch and a power supply which can be turned on or off, and alarm circuit means for producing an alarm signal when said power supply is turned off if said reset switch is not actuated.

49. The patient monitor as set forth in claim 48, wherein said alarm circuit means further includes inhibit circuit means for eliminating said alarm signal only when said power supply is turned on and said reset switch is actuated.

50. The patient monitor as set forth in claim 48, wherein said power supply further includes low voltage alarm circuit means for producing an alarm signal when the output of said power supply falls below a predetermined limit.

51. The patient monitor as set forth in claim 48, wherein said alarm circuit means includes means for inhibiting said alarm signal if said reset switch is actuated when said power supply is turned off.

52. The patient monitor as set forth in claim 33 and further in combination with a test signal generation means for producing a simulated respiration wave form, said test signal generation means having output terminals for connection to said patient unit.

53. The combination as set forth in claim 52, wherein said test signal generation means comprises an oscillator, a fixed resistance, and a semiconductor component having an input connected to an output of said oscillator and having output terminals connected across said fixed resitance.