HEART ACTIVITY INDICATOR

Abstract: Heart activity indicator, particularly a heartbeat or pulse indicator, comprising at least two electrodes (2) intended to be attached to the patient for detecting the potentials generated by the heart and means (1-12) for signal processing thereof. The heart activity indicator is provided with a monitor circuit for monitoring the contact of the electrodes (2) against the patient simultaneously with the indication of the heart activity. Said monitor circuit comprises a generator (15) for generating at least one measurement signal over the electrodes (2) and means (16) for detecting the signals which are brought back from the electrodes and comparing these signals with one or more reference signals for indication of a deficient electrode contact when a predetermined deviation between the brought back measurement signal or signals and the reference signal or signals is obtained.

Description: Heart activity indicator The present invention relates to a heart activity indicator, particularly a heartbeat or pulse indicator, having improved reliability and function and which permits a quick and simple determination of a person's heart activity or pulse even under difficult conditions. Heartbeat or pulse indicators are generally small,, portable instruments based upon detection of the potentials generated in the heart, i.e. according to the same principle as a conventional ECG apparatus. Usually two or three electrodes are used. The heartbeat indicators have found use in widely different areas but are primarily intended for use in ambulances and for observing heart patients. In order that any measurement at all could be made there must, of course, be a satisfactory contact between the electrodes and the patient's skin. To be able to determine with certainty that there is a good contact is thus directly crucial for the reliability of the measurement. In case of, for example, a heart failure, it is important that the resuscitative attempts are started as soon as possible, while resuscitative measures on a person having heart activity may instead be detrimental. Similarly, when observing a patient it is important to quickly find out if an electrode loses contact or the contact is heavily reduced. It would therefore be desirable to be able to determine by a direct measurement of the contact impedance whether there is a satisfactory contact or not. There is no such possibility in the currently existing heartbeat indicators. According to the invention there is therefore suggested a device for -detecting heart activity - hereinafter called heart activity indicator - which simultaneously with a pulse or heart activity determination going on permits the contact impedance at the electrodes to be continuously monitored. In the following heart activity indicator will mean, in addition to heartbeat or pulse indicators, also other devices for transferring or recording ECG signals. A bad or failing contact is preferably indicated through some form of alarm. For this impedance measurement the heart activity indicator of the invention is provided with a contact monitoring circuit, comprising a generator for generating at least one measurement signal over the electrodes and means for detecting the voltages generated over the electrodes and comparing these voltages with a reference voltage, which, e.g., may be the measurement signal. The monitoring circuit is preferably arranged such that signals giving the same potential at the electrode inputs are separated from signals giving different potentials at the inputs and are extracted at separate outputs. Since the heart signals, in contrast to the measurement signals, give different potentials at the electrode inputs when the electrodes are properly applied, the detection means can be caused to be affected essentially only by the measurement signals, while the latter do not O affect the measurement of the heart activity. The determination of the electrode contact according to the invention will further be completely independent of possible polarization potentials at the skin-electrode junction. The pulse measurement of a heartbeat indicator is generally arranged such that the number of heartbeats of an ECG signal is measured during a certain time and the result then extrapolated to, for example, the number of beats per minute. It is, of course, desirable to be able to determine the pulse in as short a time as possible without any substantial detriment to the accuracy of the measurement. A preferred embodiment of the invention therefore comprises a circuit with which it is possible to provide a reliable pulse measurement in only a few seconds even at a low heart activity and which is substantially insensitive to interferences in the ECG signal. Said circuit also permits an alarm for too high or low a pulse to be arranged in a simple manner. Such a circuit comprises as an essential part thereof a so-called phase locked loop locking its frequency at the average frequency of the ECG signal. For example, when used in ambulances it may some times be desirable to be able to transmit the patients ECG signals via radio to the hospital. A heartbeat indicator according to the invention is therefore suitably adapted for radio transmission of ECG signals so that it may be connected to a transmitter. By connecting a demodulator to the receiver the ECG signal may be transmitted to a recorder. In the previously mentioned pulse count the input signals are filtered such that essentially only the so-called QRS signal corresponding to the ventricle depolarization wave is recorded. In the radio transmission the whole signal complex for the heart activity should be included. On the other hand it is, of course, desired to filter out as extensively as possible the muscle interferences and other interferences which may be significant under the intended use conditions for the heartbeat indicator compared to a conventional ECG measurement in a hospital. To reduce these interferences there have previously been used low pass filters having an upper limiting frequency of 100- 150 Hz. A higher limiting frequency produces increased interferences, while a lower frequency distorts the ECG signal. A study of the power density spectrum of the ECG signal shows, however, that the major part of the power thereof is below 35 Hz. The reason that the lowest possible limiting frequency of the low pass filter still must be at such a high level is that non-phaselinear filters have hitherto been used. The phaselinear filters which are usually used in connection with time-continuous signal processing are the Bessel filters. The latter have, however, a flat slope and are thus not suitable for interference suppression. According to the invention the heart activity indicator is therefore provided with a phaselinear filter having a linear phase characteristic and a steep slope between pass band and suppression band. A suitable lower limiting frequency may be about 0,05 Hz and a corresponding upper limiting frequency about 60 Hz, particularly 5-30 Hz. Such a filter allows the ECG signal to pass undistorted, while the frequency components of the muscle interferences - which to a major part are over 30 Hz - are substantially eliminated. A suitable embodiment of such a filter is a novel realization of erner filters which have been modified according to the invention to be adapted to low frequency applications. The invention also relates to a component saving realization of such a filter which may be made with conventional operational amplifiers. In the following the heart activity indicator of the invention will be described in more detail with regard to a particular embodiment thereof, to which it, however, is not restricted in any way. Reference is made to the accompanying drawings, wherein Fig. 1 is a schematic block diagram of a heartbeat indicator of the invention; Fig. 2 is a principal diagram of an electrode contact monitoring circuit of the invention; Fig. 3 is an embodiment of the circuit of Fig. 2; Fig. 4 is a principal diagram of a phase locking circuit of the invention; Fig. 5 is an embodiment of the circuit of Fig. 4 Fig. 6 is a principal diagram of a phaselinear ilter; Fig. 7a and 7b show an embodiment of a phase-linear filter of the invention; and Fig. 8 is a perspective view of an embodiment of a heartbeat indicator of the invention. In the block diagram shown in Fig. 1 reference numeral 1 represents an input stage to which contact electrodes 2 for application to a patient are connected through leads 3. Only two contact electrodes 2 are shown in the Figure, but, of course, also three or possibly more electrodes may be used. The input stage 1 comprises a monitoring circuit for detecting the contact between the electrodes 2 and the patient. This circuit is connected to an alarm device arranged to give alarm in case of bad contact. The output from the input stage 1 is via a band pass filter 5 connected to a heart frequency converter 6 containing a phase locked loop. The latter is connected, on one hand, to a counter 7 and, on the other, to an alarm device 8 arranged to give alarm when predetermined upper and lower limiting values, respectively, of the heart frequency are passed. The counter 7 is connected to a device 9 for presenting the heart frequency, suitably in a digital form. Preferably, the output of the band pass filter 5 is also directly connected to a suitable device 10 for indicating the ventricular beats, for example a row of light emitting diodes. The device of Fig. 1 is preferably also adapted for radio transmission of the output signal from the input stage 1. To this end said output is connected or connectable to a frequency modulator 12 via a low pass filter 11. Briefly the above described device operates in the following way. The potentials generated by the heart are detected with the electrodes 2. The latter are applied to the patient in a suitable manner. When using, for example, three electrodes a suitable electrode positioning is one electrode on each wrist and the third electrode on one ankle. The voltage changes detected by means of the electrodes 2, i.e. the ECG signals, are amplified in the input stage 1. Simultaneously with the measurement of the heart signals a continuous monitoring of the' electrode contact is effected as will be described further below. In case of a bad or failing contact a suitable alarm, e.g. acoustic, is given by the alarm device . The ECG signal amplified in the input stage 1 is filtered by the band pass filter 5 before being fed to the heart frequency converter 6. The signal reaching the latter is relatively heavily filtered, such that essentially only the ventricular depolarization signal or the so-called QRS signal remains. In the heart frequency converter 6 the frequency is multiplied and counted by the counter 7 for a predetermined time, said time and multiplication being adapted such that the value presented on the display device 9 refers to, for example, the number of pulse beats during a 60 seconds interval. As will be disclosed further below the frequency converter 6 is also arranged to suppress interference signals, e.g. muscle interferences. A, for example, acoustic alarm is given by the alarm device 8 when the heartbeat frequency exceeds or falls below the preselected limiting values. At the same time as the heartbeat frequency or pulse may be read in digital form on the display means 9 each ventricular beat is indicated on the indicating device 10. For radio transmission, e.g., via an ambulance radio, the frequency modulator 12 is connected to the radio transmitter in question - as represented by the arrow 13 in the Figure. The amplified ECG signal from the input stage 1 is filtered from muscle interferences and the like through the low pass filter 11 before reaching the frequency modulator 12. A particularly advantageous such low pass filter 11 will be described in more detail below. In the following the heartbeat indicator of Fig. 1 will be described in more detail with regard to the circuit for electrode contact monitoring, the heart frequency converter circuit and the above mentioned low pass filter 11. The principle of the electrode contact monitoring circuit appears from Fig. 2, which shows a suitable design of the input stage 1. The latter contains as main components a signal separating stage 14, a generator or oscillator 15 and a comparator 16. The input of the signal separating stage l-t is connected to the electrodes 2. On the output side thereof the signal separating stage is connected, on one hand, to the blocks 5-12 in Fig. 1, represented by a heart signal indicating block 18, and, on the other hand, to the comparator 16. The signal separating 5 stage 1 is arranged such that essentially only the heart signals give an output signal at the output to the block 18, while essentially only the measurement signals give an output signal at the other output to the comparator 16. The generator 15 is arranged to feed the electrode input of the signal separating stage 14 with a voltage, the generator 15 also being arranged to provide the 10 second input of the comparator 16 with a signal. Optionally this feedback of the output signal from the generator 15 may instead be effected through a third output of the signal separating stage 14, as is indicated by the dashed connection 20. ' The signals from the generator 15 give rise to an electrode contact 15 impedance depending output signal on the output 19 to the comparator 16. The latter signal is compared with the reference signal to the second input of the comparator 16. As mentioned above this reference signal is taken directly from the generator 15 or via the signal separating stage 14. The reference signal may optionally be one of several feedback measurement signals that are brought back. 20 In case of a good contact between the electrodes and the patient the signals to the comparator 16 essentially cancel each other, while in case of a bad or failing contact an output signal from the comparator 16 is fed to the alarm device 4 such that an alarm is given. The function of the signal separating stage 14 may be based upon separation of either the frequency or the phase of the input 25 signals. In the former case the separation is effected by means of suitable filters, while in the latter case addition and subtraction of the signals are used to separate signals giving the same or different potentials, respectively, at the inputs of the separating stage 14, hereinafter called co-phasal and opposed phase signals. The measurement signals thus give co-phasal signals at the inputs, while ^ for a correct electrode application the heart signals have essentially opposed phases. Fig. 3 shows a specific example of a circuit for the input stage 1. This circuit comprises a per se conventional instrumentation amplifier built arround three operational amplifiers 21, 22 and 23. There is further an interference 35 suppression amplifier 24 for active grounding, a summer or comparator 25 and an alternating current generator 26. The shown circuit is intended to be used with three electrodes, the inputs of which are designated with A, B and C in the Figure. The non-inverting inputs of the operational amplifiers 21 and 22 are connected to the electrode inputs A and B, respectively. Optionally, each of said operational amplifier inputs are also connected to a common ground point through input resistors 27 and 28. The amplifiers 21 and 22 have a negative feedback through resistors 29, 30 and 31. The inputs of the operational amplifier 23 are coupled to the outputs of the operational amplifiers 21 and 22 through resistors 32 and 33, respectively. The operational amplifier 23 further has a negative feedback via resistors 32 and 34, and the non-inverting input thereof is grounded via a resistor 35. Further, the outputs of the operational amplifiers 21 and 22 are connected to each other through two resistors 36 and 37. The inverting input of the operational amplifier 24 is connected to a point between these two resistors, while the non-inverting input thereof is grounded via the alternating current generator 26. The output of the amplifier 24 is connected to the electrode connection C through a resistor 38. Finally the summer 25 has its inverting input connected to the non-inverting input of the operational amplifier 24 . and its non-inverting input connected to the output of the amplifier 24, optionally via a band pass filter 39. In use the two electrodes connected to the electrode inputs A and B are applied to the patient, such that the signals generated by the heart are substantially in opposite phases on the inputs A and B. This is, for example, the case . when one electrode is placed on each wrist of the patient. The third electrode, which is connected to the electrode connection C, is to serve as the ground electrode, and is placed, e.g., on one of the patient's ankles. The contact impedance of the junction between skin and electrode comprises a resistive and a capacitive component. In order that the effect of the capacϊtive component should not be predominating, the frequency of the measurement signal from the alternating current generator 26 is selected substantially in the same frequency range as the ECG signal, i.e. 0-150 Hz. From the circuit description above it appears that essentially only signals received in opposed phases on the electrode inputs A and B, respectively, give an output signal on the output D of the operational amplifier 23, while co-phasal signals are outbalanced in the latter. The measurement signal from the alternating current generator 26 which is fed to the circuit via the positive or non-inverting input of the interference suppression amplifier 24 gives rise to co- phasal signals between the electrode inputs A and B and does thus not produce any output signal at D. In case of a good contact the contact impedance at the electrodes may be neglected in comparison with the resistance over the resistors 27 and 28, respectively. Viewed from the non-inverting input of the amplifier 24 the circuit may then be regarded as a voltage follower, and the voltage of the output of the amplifier 24 will therefore be the same as the voltage of the non- inverting input thereof, i.e. the voltage U, generated by the alternating current ' generator 26. The interference suppression amplifier 24 will therefore have the same voltage on the inverting input as on the output thereof. The voltages at the input and output of the amplifier are compared in the summer 25. In case of a good electrode contact no output signal is thus obtained at the output E of the summer. The band pass filter 39 prevents 50 Hz interferences from affecting the measurement. The filter is tuned to the frequency of U. , which therefore must not be selected to be the power line frequency or multiples thereof. The ECG signals on the electrodes do not produce any output signal from the summer 25, since the voltage vectors on the inputs A and B, as mentioned above, essentially are in opposed phases when the electrodes are properly applied and therefore disappear at the summation on the non-inverting input of the interference suppression amplifier 24. On the other hand interferences due to muscle activities, which are co-phasal on the electrode inputs, may give a certain output signal from the summer 25. This is, however, an advantage, since one may hereby check that the ECG signal extracted at D is fairly free from muscle interferences. If the contact impedance at the electrodes increases through a bad or failing contact such that the effect thereof may no longer be negligible as above, the voltage at the output of the interference suppression amplifier 24 will be greater than the voltage U . generated on the input side. The voltages at the inputs of the summer 25 do therefore no longer cancel each other, and an output signal is obtained at the output E. The output voltage of the summer 25 is always zero for a good contact, while an impedance increase at any one of the electrodes produces an output signal from the summer. The output signal is detected by a not shown level comparator, which is arranged to sound an alarm if the voltage exceeds a predetermined value. While the circuit shown in Fig. 3 is adapted for three measurement electrodes, it may easily be modified for use with two electrodes. In such a case only the electrode inputs A and B are used, and the electrode connection C is connected to the input B. Optionally the interference suppression amplifier 24 may then be omitted. The above mentioned principle may, of course, also be applied to the use of more than three measurement electrodes. With the above described circuit one may thus without introducing interferences into the ECG signal continuously monitor the electrode contact while an ECG measurement is going on. As mentioned above the output signal from the operational amplifier 23 in Fig. 3 is processed by the band pass filter 5 and the frequency converter 6 before a count is performed by the counter 7 (Fig. 1). The design of the frequency converter 6 and the alarm device 8 is diagrammatically shown in Fig. 4. The central part of the circuit consists of a phase locked loop 40 connected between the output of the band pass filter 5 and the counter 7. Since the phase locked loop 40 is locked to the average frequency of the heartbeats or a multiple thereof it is possible to count the number of beats for a relatively short time and still obtain a pulse value with acceptable accuracy. To be able to give alarm at too high or low a pulse the circuit further comprises a flip-flop 41, an integrator 42 and a level comparator 43, to which the alarm device 8 is connected. The signal from the integrator 42 is also used for a rough setting of the phase locked loop 40, as will be further explained below. The circuit may optionally contain a non-linear filter 44 (indicated with dashed lines in the Figure) inserted before the input of the phase locked loop 40. The function of the non-linear filter is to change the signal such that the major part of the power is in the fundamental frequency. The flip-flop 41 may optionally be connected into the integrator loop in front of the integrator 42. An example of a frequency converter circuit according to the above is shown in Fig. 5. The phase locked loop 40 in Fig. 4 comprises in per se conventional manner a phase detector 45, the output of which is feedback-coupled to one input thereof via a loop filter and a voltage-controlled oscillator 46. In the shown case the loop filter consists of a low pass link comprised of two resistors 47a and 47b as well as a capacitor 47c. In the shown case the phase locked loop also comprises a counter 48 between the oscillator 46 and one input of the phase detector, said counter 48 operating as a divider. The other input of the phase • detector 45 is connected to the circuit input F (which is connected to the band pass filter 5 in Fig. 1) via a threshold 50, the above mentioned flip-flop 41 and a non-linear filter corresponding to the block 44 in Fig. 4. The non-linear filter is built up of two parts, viz. a first part arranged in series with the signal input of the phase detector 45, and a second part connected between the signal and feedback inputs of the phase detector. The first part comprises a diode 51 and a resistor 52 connected in parallel with a capacitor 53. The other part comprises a diode 54, a * resistor 55 connected to a voltage source G and a capacitor 56. The voltage-controlled oscillator 46 is for rough adjustment thereof connected to a current generator 57. The latter is in turn arranged to be controlled by the signal on the output of the flip-flop 41 via the integrator 42 in Fig. 4. In the shown case the latter consists of a resistor 58a and a capacitor 58b. The output of the integrating circuit is as in Fig. 4 coupled to the level comparator 43 and the alarm device 8. The above described phase locked loop is connected to the counter 7 at a point between the oscillator 46 and the counter 48. Through the threshold 50, e.g. a Schmitt trigger, the entering signal is - converted into a square wave. By suitably selecting the levels of the threshold a certain interference suppression is obtained. The levels are adapted such t only the QRS part of the ECG signal passes. The flip-flop 41 is in the shown c suitably a re-triggable monostable flip-flop, whose pulse length is adapted to t highest number of heartbeats per minute which is to be counted. Thus, if it desired to determine up to 300 beats per minute the pulse length will be seconds. Pulse beats within 0,2 seconds give rise to a longer pulse from the fl flop. The output signal from the flip-flop passes to the non-linear filter 51-56. the latter the signal is changed such that the major part of the power is in t fundamental frequency. At high frequencies the spectrum may be unchanged, b for low frequencies the bandwidth of the signal is reduced by the pulses bei extended, since too great a bandwidth of the input signal may cause the pha locked loop 45-48 to lock at overtones for a low heart frequency. The pha detector 45 may, for example, consist of an AND or OR gate but is in the sho case preferably an EXCLUSIVE OR gate. The resistor 52 is to maintain the inp of the phase detector 45 at zero when the diodes 51 and 54 block. The resistor will, however, discharge the capacitor 53, and the time constant must there o be sufficiently long for locking to take place also at low heart frequencies. At higher heart frequency the capacitor 53 must be discharged faster. This will th be effected by a negative slope of the output signal from the counter 48 via t filter part 54, 55, 56. The output signal from the phase detector 45 sets, via t loop filter 47a, 47b and 47c, the frequency of the voltage-controlled oscillat 46, said loop filter 47a, 47b, 47c being optimized with regard to a quick locki and good interference suppression. To insure locking within the whole of t actual frequency range the frequency of the oscillator 46 is roughly adjusted wi the current generator 57. The latter is controlled by the voltage from the fli flop 41 via the integrating net 58a, 58b. This voltage will be approximate proportional to the heartbeat frequency and has a long time constant, so that will only to a small extent be affected by interferences. The voltage is sensed the level comparator 43, an alarm being given by the alarm device 8 when pres limits for high and low heart frequencies, respectively, are passed. In t voltage-controlled oscillator 46 the output signal from the phase detector 45 multiplied to a suitable frequency to be counted by the counter 7 and is th divided down by a corresponding factor in the counter 48, whereupon the signal compared with the heart frequency in the phase detector 45. If thus the count 7 counts the number of periods in the output signal of the oscillator 46 during 7, seconds and the frequency is eight times higher than the heartbeat frequenc the number of beats per minute may be presented on the display device 9 in Fi 1. Through the phase locked loop 45-48 the output signal from the voltag controlled oscillator 46 is locked to an average frequency of the ECG signal, th influence of instantaneous anomalies and interferencies of the signal thereby being eliminated. Consequently a very reliable extrapolated pulse value may be obtained despite measuring for only a short period of time. When ,1 or example, presenting the heartbeat frequency digitally, the last Figure of the presentation wiil hereby be stable, so that an irritating moment in the reading is avoided. The above described circuit is also extraordinarily component saving. Thus, only two IC capsules are included in the actual loop. As mentioned above the low pass filter 11 is, for limitation of muscle interferences and other interferences of the ECG signal, preferably a phase- linear filter, i.e. a filter whose phase shift varies substantially linearly with the frequency. With such a filter the upper limiting frequency of the low pass filter may be lowered to about 45 Hz without distorting the ECG signal. It is in effect possible to use a lower limiting frequency of about 35 Hz or possibly even 30 Hz without any significant distortion of the signal. Since the major part of the frequency components of the muscle interferences is over 30 Hz, un ECG signal is hereby obtained, in which the major part of the muscle interferences has been eliminated. A particularly suitable type of filters is so-called Lerner filters, which are phase-linear filters having steep flanks. Such filters are described, e.g., in Lerner, Robert N., Proceedings of the IEEE, March 1964, p e 249-268. The design of the latter is diagramatically shown in Fig. 6. As appears from the shown circuit the input signal U. is distributed between two branches each comprising a capacitor C. and C-, respectively, and a number of series " resonance circuits Z Q , Z-, Z^ .... Z , and Z,, Z-, Z 5 .... Z , respectively, which are connected in series. The two filter branches are each connected to an input of a summer 59. The output signal from the filter U . comprises frequencies within a sharply defined interval. In Fig. 7a and 7b a modified Lerner filter is shown which is particularly adapted to the actual low-frequency application. The illustrated circuit, which has not been described earlier, is component-saving and may be realized with conventional operational amplifiers. As in Fig. 6 the filter comprises two branches each connected to an input of a summer 59. The circuit comprises three series resonance circuits 60, 61 and 62 corresponding to Z~, Z. and Z- in Fig. 6. Z Q has its correspondance in a resistor 63 connected via a voltage generator 64. C, and C 2 in Fig. 6 correspond to capacitors 65 and 66 in Fig. 7. The two filter branches are interconnected via two diodes 67 and 68 connected in parallel and serving to quickly recharge the capacitors 65 and 66 when great changes of the direct current level of the input signal U. occur. The summer 59 consists of a differential amplifier circuit comprising an operational amplifier 69, wherein the output voltage from the two filter branches are summed Yia input resistors 70 and 71. The operational amplifier 69 has a negative feedback via a resistor net 71, 72, 73 and 74, which forms a potential divider for the feedback voltage. Optionally a capacitor 75 is connected in series with the resistor 74 to increase the direct current stability of the differential stage. The non-inverting output of the operational amplifier 69 is further connected to a voltage source K via a potential divider formed by three resistors 76, 77 and 78. The construction of the series resonance circuits 60, 61 and 62, which principally are built up by a so-called D-member D and a resistor R, is shown in Fig. 7b. In contrast to the resonance circuits in a conventional Lerner filter they are based upon operational amplifiers. Each resonance circuit comprises an operational amplifier 79 with a negative feedback via a resistor 80. A resistor 81 is connected in series to the inverting input of the amplifier 79, and a resistor 82 is placed in parallel with the amplifier inputs. Further, the amplifier output is feedback-coupled to the non-inverting input via a resistor 83 and a capacitor 84. Finally a capacitor 85 is arranged in series with the capacitor 84. The. above described circuit is properly a band pass filter, the limiting frequencies of which may be selected to be, for example, 0,07 and 40 Hz, respectively. Like a conventional Lerner filter it has a linear phase characteristic and a steep slope between pass band and barrier band. At the lower limiting frequency the filter operates as an ordinary high pass RC-Iink, while it at higher frequencies gradually is transformed into a conventional Lerner filter. The specially shown filter design has seven poles, two of which are correction poles. Fig. 8 shows an example of the external design of a heartbeat indicator of the invention. The previously described circuits are disposed within a casing 86. On the face side of the casing a display device 2,7 for the pulse frequency (corresponding to 9 in Fig. 1) is provided. The latter may, for example, be a liquid crystal display indicating the number of heartbeats per minute. Further, a row of light emitting diodes 88 (corresponding to 10 in Fig. 1) is provided in the casing 86. This diode row, which e.g. may contain ten segments, is arranged to light on each ventricular beat, the reading (i.e. the number of segments which light) being proportional to the strength of the heart signal, i.e. the depolarization potential of the heart. The electrodes, which are applied to the patient via wires of a sujtable length, are connected to contacts 89. A contact 90 is arranged for battery charge and for connection to a radio transmitter. 91 indicates a combination button intended, on one hand, for battery control, and on the other hand, for the connection of a reference signal, which is necessary to indicate the ECG scale in radio transmission. Switching on and off the apparatus is effected with a control 92. The apparatus is suitably given such a size that it may easily be brought along by, e.g., ambulance men and be carried in a coat pocket or the like. The above described heartbeat indicator may be used within medical as well as preventive care. Thus, in addition to the above mentioned use in ambulances and for monitoring purposes, it may also be used in, e.g., physical form tests, in intensive care and rehabilitation and in catastrophy equipments. The invention is, of course, not restricted to the devices and circuits particularly described above and shown in the drawings, but many modifications and variations are possible within the scope of the subsequent claims. While the device and circuits according to the invention are particularly advantageous for portable apparatuses, they may, of course, also be used in stationary equipments. It is understood that the described and shown circuits may be realized as IC circuits, e.g. of a hybrid or monolithic design. ; * "- * C F

Claims: CLAIMS 1. A heart activity indicator comprising at least two electrodes (2) for application to a patient for detecting the potentials generated by the heart and circuit means (1-12) for signal processing thereof and indication of the patient's heart activity, characterized by comprising in combination with said circuit means a monitor circuit for monitoring the contact of the electrodes (2) against the patient, which circuit comprises a generator (15, 26) for generating at least one measurement signal over the electrodes (2), a separating stage (14) arranged to separate the brought back measurement signals from the heart signals in the signals detected by the electrodes (2), and comparator means (16) arranged to compare the brought back measurement signals with at least one reference signal as well as means for indicating a deficient electrode contact, when a predetermined deviation between the brought back measurement signal or signals and the reference signal or signals is obtained, said stage (14) for separating the heart signals and the measurement signals being arranged such that essentially only the heart signals give an output signal at one or more first outputs connected to said circuit means for indication of the heart activity, and that essentially only the measurement signals give an output signal at one or more second outputs sensed by said comparator means (16), whereby a measurement of the contact impedance is permitted while a heart activity indication is going on. 2. A device according to claim 1, characterized in that the stage (14) for separating the heart signals and the measurement signals is arranged such that only signals giving different potentials at the inputs of the separating stage (14) produce an output signal at said first outputs, and that signals giving the same potential at the inputs constitute at least the major part of the output signal at said second outputs. 3. A device according to claim 2, characterized in that the monitor circuit (1) comprises a first amplifier (21, 22, 23) connected to at least two electrodes (2) and having one output for the sum of the input signals and one output for the difference between the input signals, and a second amplifier (24) having an input connected to the summation signal output of the first amplifier (21, 22, 23) and having its output connected to one electrode (2), said measurement signal generator (26) being connected to a second input of said second amplifier (24), and that the comparator means (16) comprise a comparator (25) arranged to compare the state of said last-mentioned input of the second amplifier (24) with the state of the output thereof. 4. A device according to claim 1, 2 or 3, characterized in that the circuit means (1-12) for signal processing of the generated heart potentials comprises a circuit (6-9) for determining the number of heartbeats during a predetermined time, which circuit comprises a phase locked loop (40) arranged to maintain the signal intended for the heartbeat determination at the fundamental frequency of the heart signal. 5. A device according to any one of the preceding claims, characterized in that it for providing an interference suppressed ECG signal comprises a substantially phase-linear, time-continuous low pass filter (11) having an upper limiting frequency below about 60 Hz, particularly at 45-30 Hz.

Inventors: OEBERG TOMMY, OEBERG, TOMMY
Applicants: HOEOEG ANDERS, SVENSSON MATS, HOEOEG, ANDERS, SVENSSON, MATS
Application year: 1981
Main classification: A61B5
IPC codes: A61B5