5,357,969
`[11] Patent Number:
`[19]
`Unlted States Patent
`
`Herleikson
`[45] Date of Patent:
`Oct. 25, 1994
`
`l|||||||||||||Illlllll||I||||I||IIIIIllIlllIlIlI||I|||I||||||||||||||l|||Il
`USOOS3S7969A
`
`[54] METHOD AND APPARATUS FOR
`N ECG
`31%;?ARLATELY DISPLAYING A
`‘
`Inventor:
`Earl C. Herleikson, Groton, Mass.
`
`[75]
`
`4,494,551
`l/1985 Little et al.
`........................ 128/696
`Primary Examiner—William E. Kamm
`Attorney, Agent, or Firm—Brent F. Logan; Curtis G.
`Rose
`
`[73] Assignee: Hewlett-Packard Company, Palo
`Alto, Calif.
`
`ABSTRACT
`[57]
`‘
`Method and apparatus for removmg baseline wander
`from an ECG signal. The ECG signal is filtered with a
`[21] Appl. NC»2 32,895
`high-pass filter having a variable comer frequency. In
`[22] Filed:
`Mar. 18, 1993
`[51]
`Int Cl 5
`response to finding low-frequency components in the
`output of the high-pass filter, its corner frequency is
` [52] US. Cl. ........
`temporarily Increased. The corner frequency may de-
`[58] Field of Seare
`crease accordlng to a decay functlon or sensing the
`[56]
`References Cited
`absence of a low-frequency component in the filter’s
`us PATENT DOCUMENTS
`output. The corner frequency may be decreased in re-
`sponse to sensing ECG activity.
`
`3,569,852
`3/1971 Berkovits ................. 128/696
`
`4,381,786
`5/1983 Duggan ......
`128/696
`4,408,615 10/1983 Grossman ................ 128/696
`
`8 Claims, 6 Drawing Sheets
`
`CLIP TO \64
`
`50 w
`0025 Hz
`MONITOR DATA 0U
`VARIABLE
`VARIABLE
`TPUT
`
`HPF
`P60
`58
`VARIABLE
`PEAK DET.
`
`
`ATTEN
`W/DELAY
`
`
`
`66
`SLOPE
`THRESHOLD
`THRESHOLD/
`CLIP T70
`
` ACTIVITY
`
`DETECT
`DETECTACTIVITY
`<=1
`CUP T0
`0025 HZ
`
`
`
`DETECT
`
`
`
`
`62
`
`VARIABLE
`ATTEN.
`
`ECG DATA IN
`
`40H:
`
`7ISIJHZ
`LPF
`
`DIAGNOSTIC DATA OUTPUT
`VARIABLE8
`VARIABLE7
`HPF HPF 90
`HPHz ”Ax ("raisej
`848
`
`ABS
`
`91
`
`PEAK
`
`DETECT
`
`1
`
`L|FECOR454-1009
`
`1
`
`LIFECOR454-1009
`
`

`

`US. Patent
`
`Oct. 25, 1994
`
`Sheet 1 of 6
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`5,357,969
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`
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`2
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`

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`US. Patent
`
`Oct. 25, 1994
`
`Sheet 2 of 6
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`US. Patent
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`Oct. 25, 1994
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`5,357,969
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`US. Patent
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`Oct. 25, 1994
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`US. Patent
`
`Oct. 25, 1994
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`Sheet 5 of 6
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`5,357,969
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`US. Patent
`
`Oct.- 25, 1994
`
`Sheet 6 of 6
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`5,357,969
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`

`

`METHOD AND APPARATUS FOR ACCURATELY
`DISPLAYING AN ECG SIGNAL
`
`BACKGROUND OF THE INVENTION
`
`5
`
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`An electrocardiogram (ECG) measurement consists
`of measuring a small, approximately 5 milli-Volt, signal
`produced by the heart superimposed upon a relatively
`large, approximately 300 mV, low-frequency potential
`produced by the skin—to—electrode interface. The large
`low-frequency potential is called baseline wander. It is
`desirable to remove the baseline wander without alter-
`ing the ECG signal.
`The American Association of Medical Instrumenta-
`tion (AAMI) has specified two cases for the removal of 15
`the electrode offset. The first specification is for the
`case of ECG monitors. In this case, it is more important
`for the ECG signal to remain visible on the screen than
`to make a diagnosis based on the precise measurements
`of the ECG waveform. Thus, the monitor specification
`requires the frequency response to be only as low as 0.5
`Hz. The most typical filter for an ECG monitor is a
`single pole high pass filter having a 3 dB corner of 0.5
`Hz.
`The second specification is for the case of diagnostic
`ECG measurement. This specification requires a pass-
`band down to 0.05 Hz so the high pass filter causes only
`minimal distortion of the ECG signal. This maintains a
`high degree of accuracy allowing for the diagnosis of a
`heart.
`Typically, the distortion that is created by a single-
`pole high—pass filter is due to its nonlinear time delay. A
`single-pole 0.5 Hz high-pass filter can be greatly im-
`proved by giving it constant delay. Alternatively, the
`corner frequency can be decreased, increasing the fil-
`ter’s susceptibility to baseline wander.
`The effect of nonlinear time-delay distortion on the
`diagnosis of an ECG signal is most pronounced with
`respect to a calculation of ST segment elevation or
`depression. This calculation is based on the difference in
`voltage from the PQ segment just before the QRS and
`voltage of the ST segment just after the QRS. Solutions
`to this problem have in the past consisted of either de-
`creasing the frequency of the high pass filter as shown
`by the AAMI specification of diagnostic ECG instru-
`mentation for 0.05 Hz, or adding a minimum of 2 sec-
`onds of delay in order to perform a constant delay 0.5
`Hz high pass.
`For the foregoing reasons, there is a need for an ECG
`high-pass filter which has decreased susceptibility to
`baseline wander, yet provides an accurate representa-
`tion of the input ECG signal for monitor and diagnostic
`purposes.
`SUMMARY OF THE INVENTION
`
`20
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`The present invention is directed to method and ap-
`paratus which dynamically varies the corner frequency
`of an ECG high-pass filter, allowing it to increase and
`thus effectively minimize the effects of baseline wander,
`yet detect periods of ECG activity and,
`in response,
`decrease for maximum accuracy, thereby satisfying this
`need.
`In order to have a real-time display in a defibrillator-
`monitor application yet also minimize the distortion, the
`invention comprises a variable high pass filter. Two
`controls vary the frequency of the high pass filter. One
`control responds to the activity of the ECG signals to
`dramatically lower the frequency of the filter during a
`
`60
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`5,357,969
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`2
`QRS event in the ECG. By not responding to the QRS
`part of the waveform, the “tail” produced by the energy
`of the QRS complex is virtually eliminated, yet good
`removal of the baseline wander is achieved during the
`rest of the waveform. The other control responds to the
`DC offset of the output of the filter. By responding to
`the DC level at the output of the filter, the bandwidth of
`the filter can be reduced when there is very little base-
`line wander and the signal is very stable. This further
`reduces the possibility of distortion on the other parts of
`the waveform such as the T waves, which are broader
`and are not high enough in activity to shut down the
`bandwidth.
`These two methods for dynamically varying the fre-
`quency of the high pass filter, provide the best solution
`for keeping the ECG trace on screen for the operator to
`see with minimal processing delay and minimal ECG
`distortion.
`
`BRIEF DESCRIPTION OF THE DRAWINGS
`
`FIG. 1 shows a perspective view of a defibrillator-
`monitor.
`FIG. 2 shows a general block diagram of a defibrilla-
`tor including the present invention.
`FIG. 3 shows a detailed block diagram of a portion of
`the general block diagram of FIG. 2.
`FIG. 4 shows a flow chart of a peak detect having
`decay.
`FIG. 5 shows a flow chart of a second peak detect.
`FIG. 6 shows a flow chart of a method for determin—
`ing an accumulated amount w.
`DETAILED DESCRIPTION OF THE
`INVENTION
`
`Referring now to the drawings, FIG. 1 shows a defib—
`rillator 10. The defibrillator delivers an electrical im-
`pulse to a patient via cables 12 and paddles (not shown).
`The defibrillator 10 has a switch 14 for selecting the
`amount of energy to be delivered to the patient.
`Switches for initiating the discharge are typically lo-
`cated on the paddles.
`The defibrillator 10 has a display 16 for showing the
`patient’s heart waveform, enabling the operator to diag—
`nose the patient’s condition. Gain switches 18a, 18b
`enable the operator to increase or decrease the vertical
`size of the heart waveform on the display 16. A heart
`rate display 20 shows the patient’s current heart rate.
`The defibrillator 10 also has a strip recorder 22 for
`printing ECG waveforms in permanent form on paper
`strips 24.
`Referring now to FIG. 2, the defibrillator 10 provides
`the ECG signals received by the paddles and transmit-
`ted on the cables 12 to an analog to digital (A/D) con-
`verter 30. Alternatively, the ECG signals may come
`from a standard patient lead set acquire by separate
`electrodes. The output of the A/D converter is pro-
`vided to a digital signal processor 32 which filters the
`digital ECG signals and provides them to a central
`processing unit 34.
`The central processing unit 34 displays an ECG rep-
`resentation of the ECG data on the ECG display 16 and
`displays the patient’s heart rate on the heart rate display
`20. The central processing unit also controls the defib-
`rillator electronics 36.
`The central processing unit 34 also accepts input from
`the user switches 38,
`including the energy selection
`
`8
`
`

`

`3
`switch 14, the discharge switches (not shown) and the
`gain switches 18a and 18b.
`The digital signal processor 32 performs many func-
`tions on the digitized ECG signals, including low—pass
`and high—pass filtering, slope detection, activity detec-
`tion, peak detection, and attenuation. Basic to the pres-
`ent invention is the digital signal processor’s function of
`providing a high—pass variable-corner-frequency filter.
`The basic building block of a real—time variable ECG
`high-pass filter is a single-pole digital filter.
`The output y[O] at time t=0 of a digital high-pass
`filter is equal to its input x[0] at the same time less an
`accumulated amount w[O] which tracks the DC offset.
`
`5
`
`10
`
`y[0l=x[0]—W[0]
`
`(1)
`
`15
`
`The accumulated amount w[0] is equal to the accu-
`mulated amount at the previous time w[— 1] plus some
`fraction “a” of the previous difference between the
`accumulated amount and the input. That is,
`
`20
`
`Wi01=W[- 1]+a(x[~ ll-Wi-ll)
`
`which simplifies to
`
`W[Ol = W[—1]+ay[~1l-
`
`(2)
`
`25
`
`(3)
`
`The fraction, or coefficient, “a” determines the fre-
`quency response or time constant 7' of the filter accord-
`ing to the equation:
` r:
`1
`a)“,
`
`(4)
`
`where F; is the sampling frequency. For example, if
`F5: 1000 Hz, and a = 1/1000, then the time constant t
`would be 1 second and the 3 dB frequency Fe (in Hertz)
`of the filter would be
`
`
`1
`F: = 2717 '
`
`(5)
`
`By dynamically changing coefficient “a”, the respon-
`siveness of the filter can be changed.
`A single—pole low-pass digital filter is very similar to
`the high—pass filter just described. Its output is the accu-
`mulated amount w[D], rather than y[0].
`Referring now to FIG. 3, a detailed block diagram of
`a real-time variable ECG high-pass filter can be seen.
`Sampled ECG data is provided to a 40 H2 low—pass
`filter 50, a 150 Hz low-pass filter 76, and a slope detec-
`tor 66. In an exemplary version of the invention, the
`ECG data is sampled once each millisecond at 16-bit
`resolution. The 150 Hz low-pass filter 76 and the slope
`detector 66 will be discussed in more detail below.
`The 40 Hz filter defines the upper end of the passband
`for the monitor data output and is a multiple term finite-
`impulse—response (FIR) digital filter.
`The output of the 40 Hz low.pass filter 50 is provided
`to two variable high-pass filters 52 and 54 connected in
`series. The output of the second variable high pass filter
`54 is provided to the operator on the display 16. Both of
`the variable high-pass filters are single-pole high—pass
`digital filters and operate as described above, varying
`their respective comer frequencies according to a sup-
`plied coefficient “a.” The manner in which the coeffici-
`ent “a” is determined will be discussed below.
`Although a single variable high—pass filter provides
`many benefits, it fails to solve a common problem faced
`
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`4
`by defibrillators. Immediately after a defibrillation dis-
`charge, the skin—electrode interface acquires a potential
`which is exponentially dissipated, taking about 5 to 10
`seconds. For the sake of discussion, assume a linear
`decay of this voltage. The output of a single-pole high-
`pass filter approximates the derivative of its input. The
`derivative of a ramp is a slowly decaying value; thus the
`output of a single-pole high—pass filter receiving a lin-
`early dissipating voltage is a slowly decaying DC value.
`Thus, for 5 to 10 seconds after a defibrillation dis-
`charge, the patient’s ECG waveform is superimposed
`upon a steady DC value. For high decay rates, the ECG
`waveform will be off-screen.
`This DC value is typically substantial enough to ne-
`gate the benefit of the variable high-pass filter. By pro-
`viding two single—pole high-pass filters, the second filter
`will receive the DC output of the first and be able to
`eliminate the DC offset.
`The coefficient “a” supplied to the two variable high
`pass filters 52 and 54 is varied according to the DC
`offset output by the first variable high-pass filter 52. A
`higher DC offset results in a higher coefficient “a”,
`increasing the corner frequency of the high-pass filters
`52 and 54, allowing them to more quickly respond to
`the DC offset. As the DC offset decreases to zero, coef-
`ficient “a” can also effectively decrease to near zero,
`providing a highly accurate filter for stable ECG data
`signals.
`However, only the output of the first high-pass filter
`52 is used. To prevent the output of the second high-
`pass filter 54 from being away from the baseline when
`the output of the first high-pass filter 52 allows “a” to be
`reduced, the corner frequency of the second variable
`high-pass filter 54 is set to twice that of the first high-
`pass filter 52. This allows the output of the second high-
`pass filter to decay faster and thus be closer to the base-
`line than the output of the first high—pass filter 52.
`The output of the first variable high-pass filter 52 is
`supplied to a 0.25 Hz low-pass filter 56, which provides
`an output essentially equal to the DC offset from the
`first variable high-pass filter 52. This filter is a single-
`pole low-pass digital filter as described above.
`The absolute value 58 of the output from block 56 is
`supplied to a peak detector 60 having a delay. The peak
`detect allows the coefficient “a” to rise rapidly, and
`thus respond to large DC offsets quickly.
`If the coefficient “a” were allowed to decrease as
`rapidly as the DC offset from the first variable high-pass
`filter 52, then as the DC offset neared zero, “a" would
`also near zero, effectively prolonging the ultimate elimi-
`nation of the last bit of DC offset. Thus, the decay after
`a delay prohibits the coefficient “a” from decreasing
`until after the DC offset has reached zero.
`The peak detector 60 uses two storage registers:
`“peak" and “decay.” “Peak” stores a value representing
`the current peak in input values. “Decay” stores a value
`less than one which gets multiplied by the value in
`“peak,” resulting in the output of the peak detector. The
`value in “decay” is decreased periodically to slowly
`decrease the output.
`The output of the peak detector 60 is scaled such that,
`if it were passed directly to the first variable high-pass
`filter 52 as coefficient “a”, a one milIi-Volt output from
`the 0.25 low-pass filter 56 would result in a corner fre-
`quency of 0.25 Hz.
`Referring now to FIG. 4, every five milliseconds, the
`peak detector 60 gets 100 the next absolute value 58 of
`
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`

`5,357,969
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`6
`An activity detector 68 receives the slope Z from the
`slope detector 66 and provides as an output a signal
`which is the average of the absolute values of the last
`four slopes.
`
`1’10] = lZIOII +|Z[—1l| +4|Z[—2]l + 121-3]!
`
`(9)
`
`The output of the slope detector 68 cannot be used
`directly because its output can drop near zero when the
`slope of the input ECG signal changes signs. This will
`occur at the peak of each R wave. Other equations for
`an activity detector can be used. The requirements are
`that the activity Y[0] remain high for periods of in-
`creased activity in the ECG data.
`A threshold detector 70 low-pass filters the output of
`the activity detector 68 with a corner frequency of 0.1
`Hz. Thus, the output of the threshold detector is the
`near-DC component of the slope of the input ECG data,
`and serves as a threshold for altering the coefficient “a.”
`The threshold detector is a single-pole low-pass digital
`filter with a corner frequency of 0.1 Hz, implemented as
`described above.
`Block 72 takes the ratio of threshold to activity, and
`supplies it to a clipper 74. If the ratio is greater than one,
`then the activity is less than the threshold, and no modi-
`fication of “a” will occur as a result of the ECG activ-
`ity.
`However, if the ratio is less than one, then the activity
`is greater than the threshold and the ECG signal is in a
`period of increased activity. Thus, the passband of the
`variable highspass filters 52, 54 should be decreased to
`decrease the QRS signal’s effect on the filter’s accumu-
`lated amounts w.
`
`The variable attenuator takes the output of the peak
`detector 60 and multiplies it by the output of the clipper
`74. Above, to keep the discussion of varying the coeffi-
`cient “a” based on the DC offset of the filters 52 and 54
`simple, the variable attenuator was described as passing
`the output of the peak detector 60 directly to the clipper
`64. During periods of low ECG activity, the output of
`clipper 74 will be one, and the simplified description is
`correct.
`
`However, during times of high ECG activity, that is,
`the output of clipper 74 is less than one, the output of
`the variable attenuator will be decreased by the ratio of
`threshold to activity as provided by block 72. This has
`the effect of decreasing the coefficient “a” supplied to
`the first and second variable high-pass filters 52 and 54
`during times of increased ECG activity.
`The 150 Hz low-pass filter 76 mentioned above de—
`fines the upper end of the passband for the diagnostic
`data output and is a multiple term finite-impulse-
`response (FIR) digital filter.
`The output of the 150 Hz low-pass filter 76 is pro-
`vided to a third and fourth variable high-pass filters 78
`and 80 connected in series. The diagnostic output of the
`fourth variable high-pass filter 80 is provided to the
`user. Thus, their primary purpose is to maintain the
`ECG signal with the lowest possible corner frequency
`of high-pass filtering.
`The third and fourth variable high-pass filters 78, 80
`are single-pole high-pass digital filters and operate as
`described above. They both operate with the same cor-
`ner frequency, unlike the first and second variable high—
`pass filters 52 and 54 discussed above in reference to the
`monitor data output.
`
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`the low pass filter 56. If 102 the value is 3.1 percent
`greater than the peak detector’s current output, then the
`value is stored 104 in the “peak” register and the value
`0.97 is stored 106 in the “decay” register. The output is
`then equal to the product of the values stored in the
`“peak” and “decay” registers 110.
`However, if 102 the value is not 3.1 percent greater
`than the peak detector’s current output, then the value
`in the “decay" register is updated 108 according to the
`following equation.
`
`decay(n + 1): decay(n)~0.0045(l m decay(rz))
`
`(6)
`
`As the value in the “decay” register decreases, its rate
`of decrease becomes greater. Equation 6 can be solved
`for decay(n) yielding:
`
`decay(n) = 1 — (1 —decay(0))(1 +0.0045)".
`
`(7)
`
`Thus, the value stored in the “decay” register, assum-
`ing no new peak is detected in block 102, will decay as
`shown in the following table.
`TABLE 1
`SECONDS
`"DECAY" REGISTER
`1.34
`0.90
`2.11
`0.30
`2.56
`0.70
`3.13
`0.50
`3.58
`0.25
`3.90
`0.00
`
`11
`253
`422
`512
`626
`716
`780
`
`Referring again to FIG. 3, the output of the peak
`detector 60 is provided to a variable attenuator 62,
`which for the present will be described as passing the
`output of the peak detector 60 on to the clipper 64. The
`description of the variable attenuator 62 will be aug-
`mented below.
`The clipper 64 provides as its output the greater of (1)
`the output of the variable attenuator 62, and (2) a coeffi-
`cient “a” corresponding to a comer frequency of 0.025
`Hz for the first variable high-pass filter 52. Thus, the
`minimum corner frequency for the first variable high-
`pass filter 52 is 0.025 H2 and for the second variable
`high-pass filter 54 is 0.050 H2.
`The coefficient “a” to the first and second variable
`high-pass filters 52, 54 can also be varied by the “activ-
`ity” of the input ECG data signal.
`A digital triangular convolution filter operates on the
`input ECG data and provides as an output a “slope” Z
`which corresponds to the average slope of the last 40 ms
`of ECG data. For this reason, the digital triangular
`convolution filter is termed herein as a “slope detector”
`66. When the ECG signal has high activity, such as
`during the QRS complex, the output of the slope detec-
`tor will be elevated. When the ECG signal is essentially
`flat, the output of the slope detector will be zero.
`The slope detector’s coefficients are selected such
`that its output slope is the average slope of the last 40 ms
`of the ECG data. An exemplary equation used by the
`slope detector 66 on ECG data sampled at 5 ms inter-
`vals is given in the following equation.
`
`Z(t)={—32X[—7]—29X[—-6]—-20X[—S]
`
`—7X[—4]+7X[—3]+20X[—2]
`
`+29X[—1]+32X[0]}/ 64
`
`(8)
`
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`7
`The output of the 150 Hz low-pass filter is also pro—
`vided to a 2 Hz single-pole high-pass digital filter 82.
`The 2 Hz corner frequency is set so that the filter’s
`output contains virtually no baseline wander.
`The output of the 2 Hz high—pass filter 82 is sub—
`tracted 84, 86 from the outputs of the third and fourth
`variable high-pass filters 78, 80. The resulting differ-
`ences are bandpass filters having passbands between 2
`Hz and the corner frequencies of the third and fourth
`variable high-pass filters 78, 80. The maximum of the
`absolute values of the two differences is provided to
`block 90 by block 88.
`Block 90 scales the output of block 88 to the gain of
`the ECG display 16 (FIG. 1). It does this by dividing
`the output of block 88 by the voltage which represents
`the extreme edge of the output device, such as the paper
`edge 24 of the strip recorder 22, and squaring the result.
`The output of block 90 is provided to a peak detector
`91. The peak detector provides an output to the third
`and fourth variable high-pass filters 78, 80 through a
`second variable attenuator 92 and clipper 94.
`Referring now to FIG. 5, the peak detector 91 uses a
`storage register “peak” for storing peak values detected
`in the output of block 90. As a first step, the peak detec-
`tor 91 gets 150 the output of block 90.
`That output is scaled 152 such that an output of unity
`from block 90 would result in a corner frequency of
`0.025 Hz at variable high-pass filters 78, 80. If 154 the
`resulting scaled value “tmp” is greater than the value
`stored in “peak,” then the value in “trnp” is stored in
`“peak” and provided 164 as the output of the peak
`detector 91.
`However, if 154 the resulting scaled value “tmp” is
`less than the values stored in “peak,” then the output of
`block 90 is scaled 158 such that an output of unity from
`block 90 would result in a corner frequency of 0. 10 Hz.
`If 160 the resulting scaled value “tmp” is less than the
`value stored in “peak,” then the value in “tmp” is stored
`in “peak” and provided 164 as the output of the peak
`detector 91.
`Thus, the output from the peak detector 91 changes
`in response to two different conditions. If the output
`from block 88 is so large that it would exceed the cur-
`rent peak value, then the peak detector’s output changes
`to increase the comer frequency of the variable high-
`pass filters 78, 80. If the output from block 88 is so small
`that four times its value does not exceed the current
`peak value, then the peak detector’s output changes to
`decrease the comer frequency of the variable high-pass
`filters 78, 80.
`The second variable attenuator 92 operates similarly
`to the first variable attenuator 62. It takes the output of
`block 91 and multiplies it by the output of the clipper
`74. During periods of low ECG activity, the output of
`the clipper 74 will be one, and thus the output of the
`second variable attenuator will equal the output of
`block 91. Otherwise, the output of the second variable
`attenuator will be decreased proportionally to the ratio
`of the output of the threshold detector 70 to the output
`of the activity detector 68.
`The output of the second variable attenuator is
`clipped such that the resultant “a” supplied to the third
`and fourth variable high-pass filters 78 and 80 results in
`corner frequency of 0.025 H2.
`A further refinement in the control of coefficient “a”
`is preferred. As described above, the accumulated value
`w in a single—pole filter tracks the low-frequency com-
`ponents of the input signal x.
`
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`8
`During times of high activity in the ECG signal, the
`activity detector 68 causes rapid reductions in the coef-
`ficient “a" through the actions of the first and second
`variable attenuators 62 and 92, thereby preventing w
`from being affected by the QRS complex. This reduc-
`tion in coefficient “a” also prevents w from accurately
`tracking any low-frequency baseline wander present,
`thereby affecting the appearance of the signals at the
`outputs of the variable high-pass filters 54 and 80.
`By continuing to change the accumulated amount w
`during times of high activity at the same rate as it was
`changing just before activity occurred, the accumulated
`amount w will more accurately track baseline wander.
`This can be accomplished by varying the manner in
`which the accumulated amount is updated.
`Let “211" be the coefficient “a” after being reduced by
`QRS activity in the variable attenuators 62 and 92 and
`let “a2” be the difference between the two: a—al. Fi—
`nally, let “slope” be w[n]—w[n—l] where n represents
`the sample time at which the output of block 72 is one,
`that is, the last sample at which “a” was not diminished
`by QRS activity. Then the accumulated amount w can
`maintain a constant rate of change during times of high
`QRS activity according to the following equation.
`
`W[0]=w{-ll+a1y[—ll+(a2/a)slope
`
`(10)
`
`Referring now to FIG. 6, a flow chart of the above
`refinement is shown. The outputs of the peak detector
`60 and the clipped ratio from clipper 74 are read 120.
`Coefficient “a” is set 122 to the output of the peak de-
`tector and “a1" is set 124 to the product of the two read
`values. Value a2 is set to the difference between “a" and
`“a1.” If 128 that difference is zero, then it is not a time
`of high QRS activity and slope is updated 130. Then
`w[0] is determined according to equation (10), given
`above.
`Although the present invention has been described in
`considerable detail with reference to certain preferred
`versions and values, other versions are possible.
`The described version uses two variable high-pass
`filters in series 52, 54, and 78, 80 to eliminate the effects
`of a constant slope in the offset signal superimposed on
`the ECG signal. As a baseline wander filter according
`to the present invention has uses in devices other than a
`defibrillator, a single variable high-pass filter may be
`used.
`The described version changes the corner frequency
`of the variable high—pass filters 52, 54, 78, 80 according
`to both the DC offset of the input ECG data and the
`QRS activity. A variable high-pass filter according to
`the present invention may be built which varies its cor-
`ner frequency according to either DC offset, or QRS
`activity, or both.
`Therefore,
`the spirit and scope of the appended
`claims should not be limited to the description of the
`preferred versions contained herein.
`What is claimed is:
`1. A method for accurately displaying an ECG signal,
`comprising the steps of:
`receiving ECG input data;
`filtering said ECG input data with a high pass filter
`having a variable corner frequency, thereby creating
`an ECG signal;
`detecting the presence of a QRS event in said ECG
`input data;
`
`11
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`

`9
`in response to said detecting the presence of a QRS
`event step, decreasing said variable corner frequency
`of said high pass filter; and
`displaying said ECG signal.
`2. The method of claim 1, further comprising the
`steps of:
`detecting the presence of a DC offset caused by a defi-
`brillation discharge in said ECG input data; and
`in response to said detecting the presence of a DC offset
`step, increasing said variable corner frequency of said
`high pass filter.
`3. An apparatus for accurately displaying an ECG
`signal, comprising:
`a high pass filter having an input, an output, and a vari-
`able corner frequency, wherein the input of said high
`pass filter is for connection to ECG input data, and
`wherein the output of said high pass filter is for con-
`nection to a display device, said display device for
`displaying said ECG signal;
`first detection circuitry for detecting the presence of a
`QRS event in said ECG input data; and
`attenuation circuitry for decreasing said first variable
`corner frequency of said high pass filter in response to
`said first detection circuitry detecting the presence of
`a QRS event.
`4. The apparatus of claim 3 , further comprising:
`second detection circuitry for detecting the presence of
`a DC offset caused by a defibrillation discharge in
`said ECG input data; and
`said attenuation circuitry also for increasing said vari-
`able corner frequency of said high pass filter in re-
`sponse to said second detection circuitry detecting
`the presence of said DC offset.
`
`10
`5. An apparatus for accurately displaying an ECG
`signal, comprising:
`a first high pass filter having an input, an output, and a
`first variable corner frequency, wherein the input of
`said first high pass filter is for connection to ECG
`input data;
`a second high pass filter having an input, an output, and
`a second variable corner frequency, wherein the
`input of said second high pass filter is connected to
`the output of said first high pass filter and wherein the
`output of said second high pass filter is for connection
`to a display device, said display device for displaying
`said ECG signal;
`first detection circuitry for detecting the presence of a
`QRS event in said ECG input data; and
`attenuation circuitry for decreasing said first and second
`variable corner frequencies of said first and second
`high pass filters in response to said first detection
`circuitry detecting the presence of a QRS event.
`6. The apparatus of claim 5, further comprising:
`second detection circuitry for detecting the presence of
`a DC offset caused by a defibrillation discharge in
`said ECG input data; and
`said attenuation circuitry also for increasing said first
`and second variable comer frequencies of said first
`and second high pass filters in response to said second
`detection circuitry detecting the presence of said DC
`offset.
`7. The apparatus of claim 6, wherein said second
`variable corner frequency is larger than said first vari-
`able corner frequency.
`8. The apparatus of claim 6, wherein said second
`variable corner frequency is approximately twice as
`large as said first variable corner frequency.
`*
`*
`=0:
`*
`*
`
`5,357,969
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`5
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`12
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`12
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`