JPH06277189A - Method for cancelling noise and noise cancelling circuit in electrocardiograph - Google Patents
Method for cancelling noise and noise cancelling circuit in electrocardiographInfo
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- JPH06277189A JPH06277189A JP5068009A JP6800993A JPH06277189A JP H06277189 A JPH06277189 A JP H06277189A JP 5068009 A JP5068009 A JP 5068009A JP 6800993 A JP6800993 A JP 6800993A JP H06277189 A JPH06277189 A JP H06277189A
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- noise
- circuit
- electrocardiogram
- filter
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Abstract
Description
【０００１】[0001]
【産業上の利用分野】本発明は心電計におけるノイズキ
ヤンセル方法および心電計におけるノイズキヤンセル回
路に関し、特にデイジタルデータに変換して圧縮処理が
容易な出力が得られ、デジタル・ホルタ心電計に最適な
ノイズキヤンセル方法およびノイズキヤンセル回路に関
するものである。BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to a noise canceling method in an electrocardiograph and a noise canceling circuit in an electrocardiograph, and more particularly to a digital Holter electrocardiograph capable of obtaining an output which is easily converted by converting it into digital data. The present invention relates to an optimal noise canceling method and a noise canceling circuit.
【０００２】[0002]
【従来の技術】近年、心臓疾患の発見及びその状況を正
確に認識するため、ホルタ心電計で長時間連続して心電
波形を記録し、後に心電図解析装置などでこの記録波形
を再生して波形変化を判別する装置が登場してきてい
る。これらの装置におけるホルタ心電計は記録時間が２
４時間、又はそれ以上となるものある。2. Description of the Related Art In recent years, in order to detect a heart disease and accurately recognize its condition, an electrocardiographic waveform is continuously recorded for a long time by a Holter electrocardiograph, and then the recorded waveform is reproduced by an electrocardiogram analyzer or the like. A device for discriminating a waveform change has appeared. The Holta ECG in these devices has a recording time of 2
It can be 4 hours or more.
【０００３】従来のホルタ心電計は収集した心電図情報
をそのままのアナログ信号としてカセツト磁気テープ装
置に記録するものであつた。この従来のホルタ心電計
は、長時間の記録を可能とするため、その解像度も高く
出来ず、波形歪なども大きく、精度上の問題となつてい
た。また、可動部を有するため振動などの影響も受けや
すく、常時携帯する必要のあるホルタ心電計では信頼性
を落とす一因ともなつていた。The conventional Holter electrocardiograph records the collected electrocardiographic information as an analog signal on the cassette magnetic tape device. Since this conventional Holter electrocardiograph can record for a long time, its resolution cannot be increased and waveform distortion is large, which is a problem in accuracy. In addition, since it has a movable part, it is easily affected by vibrations and the like, which is one of the causes of reducing the reliability of the Holter electrocardiograph which must be always carried.
【０００４】以上の欠点を改良する方法としては、誘導
電極よりの心電波形をアナログ信号として記録するので
はなく、一旦デジタル信号に変換し、デジタル信号とし
て記録する方法が考えられる。すなわち、一旦デジタル
信号に変換し、カセツト磁気テープ装置に記録する変わ
りに、デジタル信号としてＤＡＴに記録する方法であ
る。As a method of improving the above-mentioned drawbacks, a method of converting the electrocardiographic waveform from the induction electrode into a digital signal and then recording the digital signal instead of recording the analog signal is considered. That is, it is a method of recording in the DAT as a digital signal instead of once converting it into a digital signal and recording it in the cassette magnetic tape device.
【０００５】しかし、ＤＡＴは高価であり、また、より
精密な記録が必要であり、可動部を有することより振動
などによる悪影響を防ぐことは出来なかつた。しかも、
消費電力も多く、動作電力を供給する電源も大型のもの
が必要であり、小型計量化にも限界があつた。以上の欠
点を解決する一手段として、デジタルデータをＤＡＴで
はなくＲＡＭに記録することも考えられる。However, DAT is expensive, more precise recording is required, and it is impossible to prevent adverse effects due to vibration and the like because it has a movable portion. Moreover,
It consumes a lot of power and needs a large power supply to supply operating power, and there is a limit to miniaturization. As one means for solving the above drawbacks, recording digital data in RAM instead of DAT can be considered.
【０００６】[0006]
【発明が解決しようとする問題点】しかしながら、必要
な精度を保つて長時間の心電図情報を記録するには非常
に大容量のメモリが必要であり、記憶できる心電図情報
の量にも限界があり、小型計量化、価格の低減化にも限
界があつた。このため、心電図データに圧縮処理を施
し、記憶すべき心電図情報の量を減らすことが望まし
い。しかしながら、心電図データには種々のノイズが含
まれており、圧縮率が上がらない問題点があつた。一
方、心電図データのノイズなどは忠実に再現する必要が
なく、冗長な情報も多く含まれている。However, a very large capacity memory is required to record the electrocardiographic information for a long time while maintaining the required accuracy, and the amount of electrocardiographic information that can be stored is limited. However, there were limits to miniaturization and weight reduction and price reduction. Therefore, it is desirable to compress the electrocardiogram data to reduce the amount of electrocardiographic information to be stored. However, the electrocardiogram data contains various noises, and there is a problem that the compression rate does not increase. On the other hand, it is not necessary to faithfully reproduce noise and the like in the electrocardiogram data, and much redundant information is included.
【０００７】[0007]
【問題点を解決するための手段】本発明は上述の問題点
を解決することを目的として成されたもので、上述の問
題点を解決する一手段として以下の構成を備える。即
ち、誘導電極よりの心電図情報に含まれるノイズ成分の
パワーを検出する検出手段と、該検出手段の検出パワー
が所定以上の場合には前記誘導電極よりの心電図情報に
対してノイズ除去フイルタを介してノイズキヤンセルを
行なつた後に基線変動除去を行ない、前記検出手段より
の検出パワーが所定以上でない場合にはノイズ除去フイ
ルタを介さずそのまま基線変動除去を行なうノイズ除去
手段とを備える。The present invention has been made for the purpose of solving the above-mentioned problems, and has the following structure as one means for solving the above-mentioned problems. That is, the detection means for detecting the power of the noise component contained in the electrocardiogram information from the induction electrode, and when the detection power of the detection means is a predetermined value or more, the noise removal filter is applied to the electrocardiogram information from the induction electrode. Noise canceling means for removing the baseline fluctuation after performing noise canceling, and if the detected power from the detecting means is not more than a predetermined value, the noise removing means for directly removing the baseline fluctuation without using the noise canceling filter.
【０００８】そして、例えば前記検出手段を少なくとも
２段よりなる偶数段の微分回路と、該微分回路よりの微
分波形の絶対値を出力する絶対値回路と、該絶対値回路
に接続されたＣＩＣフイルタ回路で構成する。Then, for example, the detecting means is an even-numbered stage differential circuit having at least two stages, an absolute value circuit for outputting the absolute value of the differential waveform from the differential circuit, and a CIC filter connected to the absolute value circuit. It consists of a circuit.
【０００９】[0009]
【作用】以上の構成において、心電図情報中のノイズを
有効にキヤンセルすることができ、その後に圧縮処理を
行なう際に十分な圧縮を行なうことができ、長時間の心
電図情報を記録する際に必要とする記憶容量を低く抑え
ることが可能となる。しかも、この検出手段を加算器を
基本とした乗算器を必要としない構成とすることによ
り、高速処理が実現し、回路構成も簡単なものと出来
る。With the above structure, noise in the electrocardiogram information can be effectively canceled, and sufficient compression can be performed in the subsequent compression process, which is necessary when recording electrocardiogram information for a long time. It is possible to keep the storage capacity to be low. Moreover, by arranging this detecting means so as not to require a multiplier based on an adder, high-speed processing can be realized and the circuit structure can be simplified.
【００１０】[0010]
【実施例】以下、図面を参照して本発明に係る一実施例
を詳細に説明する。図１は本発明に係る一実施例のノイ
ズキヤンセル回路を装備したホルタ心電計のブロツク図
である。図中、１〜３は生体よりの心電図情報を収集す
る生体誘導電極（ピツクアツプ）、４は生体誘導電極１
〜３よりの収集生体信号（心電図情報）を増幅するアン
プ回路、５はアンプ回路４よりの心電図信号よりノイズ
信号を除去する本実施例のノイズキヤンセル回路、６は
ノイズの除去された心電図信号を対応するデジタル信号
に変換し、同時に圧縮処理を施す符号化圧縮回路、７は
符号化圧縮回路で圧縮処理されたデジタル心電図情報を
ＲＡＭ８に書き込む制御を行なう書き込み回路、８は心
電図情報を記憶するＲＡＭであり、例えば着脱可能なＩ
Ｃカードで構成され、当該ＩＣカードには記憶された心
電図情報が不用意に消去されないようにバツクアツプ用
の電源が装備されていることが望ましい。このため、心
電図情報の記憶が終了したＲＡＭ（ＩＣカード）をホル
タ心電計より取り外し、不図示の心電図解析装置に装着
することにより、被検者の心電図を解析し、診断の一助
とすることができる。DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS An embodiment according to the present invention will be described in detail below with reference to the drawings. FIG. 1 is a block diagram of a Holter electrocardiograph equipped with a noise-cancell circuit according to an embodiment of the present invention. In the figure, 1 to 3 are bioinduction electrodes (pickups) for collecting electrocardiographic information from a living body, and 4 are bioinduction electrodes 1.
3 to 3 are amplifier circuits for amplifying the collected biomedical signals (electrocardiogram information), 5 is the noise cancel circuit of the present embodiment for removing noise signals from the electrocardiogram signals from the amplifier circuit 4, and 6 is the noise-free electrocardiogram signals. An encoding / compressing circuit for converting into a corresponding digital signal and simultaneously performing compression processing, a writing circuit 7 for controlling to write the digital electrocardiographic information compressed by the encoding / compressing circuit into the RAM 8, and a RAM 8 for storing electrocardiographic information. And, for example, removable I
It is preferable that the IC card is equipped with a backup power source so that the stored electrocardiogram information is not erased carelessly. Therefore, by removing the RAM (IC card) in which the electrocardiogram information has been stored from the Holter electrocardiograph and mounting it on an electrocardiogram analyzer (not shown), the electrocardiogram of the subject is analyzed to aid diagnosis. You can
【００１１】図２に図１に示す本実施例のノイズキヤン
セル回路５の詳細ブロツク構成を示す。デジタルホルタ
心電計においては、データ圧縮が必要不可欠であるが、
心電図データのノイズなどは忠実に再現する必要がな
く、冗長な情報が多く含まれている。一方、このノイズ
情報等も一緒に圧縮処理して記憶させてはデータの圧縮
率も上がらず、記憶に必要な情報量も多くなつてしま
う。しかしながら、心電図データのノイズなどは忠実に
再現する必要もない。このため、本実施例においては、
心電図情報の圧縮処理を行なう以前に、必要に応じてノ
イズ除去を行ない、圧縮率を上げる様にしている。FIG. 2 shows a detailed block configuration of the noise cancel circuit 5 of this embodiment shown in FIG. Data compression is indispensable in the digital Holter monitor,
Noise in ECG data does not need to be faithfully reproduced and contains a lot of redundant information. On the other hand, if the noise information and the like are also compressed and stored together, the data compression rate will not increase, and the amount of information required for storage will also increase. However, it is not necessary to faithfully reproduce the noise of the electrocardiogram data. Therefore, in this embodiment,
Before performing the compression processing of the electrocardiogram information, noise is removed as necessary to increase the compression rate.
【００１２】図２において、１０は位相調整回路であ
り、ノイズ除去フイルタ回路２０を通した場合と該回路
２０を通さない場合との位相を一致させるために入力波
形をノイズ除去フイルタ回路２０を通した場合と同じ時
間だけ遅延させる。２０はノイズ除去フイルタであり、
ＣＩＣ２段フイルタ（２１，２２）で構成されている。
この２段フイルタとしたのは、符号化圧縮回路６での符
号化におけるサンプリング周波数に合わせるためであ
る。すなわち、本実施例ではサンプリング周波数が１８
０ヘルツであり、偶数段とすることによりサンプリング
周波数と一致させている。なお、このＣＩＣフイルタ２
１，２２は後述するノイズ検出回路３０のＣＩＣフイル
タ３３と同様の構成と出来るIn FIG. 2, reference numeral 10 denotes a phase adjusting circuit, which passes an input waveform through the noise removing filter circuit 20 in order to match the phase when the noise removing filter circuit 20 is passed and when the circuit is not passed. Delay for the same amount of time you did. 20 is a noise removal filter,
It is composed of two-stage CIC filters (21, 22).
This two-stage filter is used in order to match the sampling frequency in the coding in the coding compression circuit 6. That is, in this embodiment, the sampling frequency is 18
The frequency is 0 Hz, and the sampling frequency is matched by setting it to an even number of stages. In addition, this CIC filter 2
1, 22 can have the same configuration as the CIC filter 33 of the noise detection circuit 30 described later.
【００１３】３０はノイズ検出回路であり、ノイズ成分
の平均パワーを検出し、当該ノイズ成分の平均パワーが
所定域値ｘ以上か否かを検出し、その結果を制御信号３
５として出力する。ノイズ検出回路３０は微分回路３
１、絶対値回路３２、ＣＩＣ回路３３、量子化回路３４
で構成されている。ノイズ検出回路３０よりの制御信号
３５は切替回路４０に入力される。A noise detection circuit 30 detects the average power of the noise component, detects whether the average power of the noise component is a predetermined threshold value x or more, and outputs the result to the control signal 3
Output as 5. The noise detection circuit 30 is a differentiation circuit 3
1, absolute value circuit 32, CIC circuit 33, quantization circuit 34
It is composed of. The control signal 35 from the noise detection circuit 30 is input to the switching circuit 40.
【００１４】切替回路４０は位相調整回路１０よりの信
号又はノイズ除去フイルタ回路２０よりの信号のいずれ
か一方の信号を選択して基線変動除去フイルタ回路５０
に出力する。すなわち、制御信号が“１”の場合（ノイ
ズ成分の平均パワーが所定閾値ｘ以上の場合）にはノイ
ズ除去フイルタ回路２０よりの信号を選択し、制御信号
３５が“０”の場合には位相調整回路１０よりの信号を
選択する。The switching circuit 40 selects either the signal from the phase adjustment circuit 10 or the signal from the noise removal filter circuit 20 to select the baseline fluctuation removal filter circuit 50.
Output to. That is, when the control signal is "1" (when the average power of the noise component is greater than or equal to the predetermined threshold value x), the signal from the noise removal filter circuit 20 is selected, and when the control signal 35 is "0", the phase is selected. The signal from the adjusting circuit 10 is selected.
【００１５】５０は基線変動除去フイルタ回路であり、
心電図データの基線変動を除去する。この基線変動除去
フイルタ５０としては、ＣＩＣ２段ノツチフイルタ等で
構成でき、例えば特願昭５７−１８９２５５号に記載の
フイルタなどを採用すればよい。この基線変動除去フイ
ルタ５０よりの出力は、基線変動が除去されているのみ
ならず、ノイズ成分も大幅に抑えられており、その後の
符号化圧縮回路６での圧縮率を上げることが可能とな
る。Reference numeral 50 is a baseline fluctuation eliminating filter circuit,
Eliminate baseline fluctuations in ECG data. The base line fluctuation removing filter 50 can be constituted by a CIC two-stage notch filter or the like. For example, the filter described in Japanese Patent Application No. 57-189255 may be adopted. The output from the baseline fluctuation removing filter 50 has not only the baseline fluctuation removed, but also the noise component greatly suppressed, and it is possible to increase the compression rate in the encoding compression circuit 6 thereafter. .
【００１６】以上におけるノイズキヤンセル回路５の入
力信号を図４の（Ａ）に、ＣＩＣフイルタ３３出力であ
るノイズ成分のパワー検出結果を図４の（Ｂ）に、ノイ
ズ検出回路３０よりの出力制御信号３５を図４の（Ｃ）
に、切替回路４０の出力波形を図４の（Ｄ）に示す。な
お、ノイズ検出回路３０の微分回路３１は、心電図情報
のノイズ成分が心電図データに比し主に高周波成分であ
ることに着目して一種のハイパスフイルタ（ＨＰＦ）を
構成するための回路である。この微分回路３１により比
較的心電図波形成分の少ない帯域でのノイズ成分を取り
出している。The input signal of the noise cancel circuit 5 is shown in FIG. 4 (A), the power detection result of the noise component output from the CIC filter 33 is shown in FIG. 4 (B), and the output control from the noise detection circuit 30 is controlled. The signal 35 is shown in FIG.
The output waveform of the switching circuit 40 is shown in FIG. The differentiating circuit 31 of the noise detecting circuit 30 is a circuit for configuring a kind of high-pass filter (HPF), focusing on the fact that the noise component of the electrocardiogram information is mainly a high frequency component as compared with the electrocardiogram data. The differentiating circuit 31 extracts the noise component in the band where the electrocardiogram waveform component is relatively small.
【００１７】絶対値回路３２では、微分回路３１よりの
ノイズ成分のパワー相当を、絶対値を取ることによりＤ
Ｃ成分に移動させている。そして、続くＣＩＣフイルタ
３３によりそのＤＣ成分（ノイズのパワー相当）を抽出
する。このフイルタ３３における心電図波形のＲ波成分
はノイズ成分に比べパワーが少ないので、このＣＩＣフ
イルタ３３を通すことで、略ノイズ成分の平均パワーの
みの出力とできる。In the absolute value circuit 32, the power equivalent to the noise component power from the differentiating circuit 31 is calculated by taking an absolute value to obtain D.
It is moved to the C component. Then, the DC component (corresponding to the power of noise) is extracted by the subsequent CIC filter 33. Since the R wave component of the electrocardiogram waveform in the filter 33 has less power than the noise component, it is possible to output only the average power of the noise component by passing through the CIC filter 33.
【００１８】その後段の量子化回路３４では、このＣＩ
Ｃフルタ３３よりのパワー信号が所定閾値ｘ以上か否か
で量子化し、タイミング調整して制御信号３５として出
力する。そして、無処理データとノイズ除去フイルタ出
力データとの切替を行なう。本実施例における量子化回
路３４での閾値の決定方法は、“ノイズデータ”を入力
とした場合のＣＩＣフイルタ３３の出力特性と、“ノイ
ズのない対象データ”を入力した場合のＣＩＣフイルタ
３３の出力特性を検討して決定する。In the subsequent quantizing circuit 34, the CI
Quantization is performed depending on whether the power signal from the C filter 33 is greater than or equal to a predetermined threshold value x, timing is adjusted, and the control signal 35 is output. Then, the unprocessed data and the noise removal filter output data are switched. The method of determining the threshold value in the quantization circuit 34 in this embodiment is as follows: the output characteristics of the CIC filter 33 when "noise data" is input and the output characteristics of the CIC filter 33 when "noise-free target data" is input. Determine by examining output characteristics.
【００１９】本実施例においては、“ノイズデータ”と
して周波数に対して一様である“ガウシアンホワイトノ
イズ”を用い、このノイズがバーストし出現するデータ
列を作り、このデータ列をノイズキヤンセル回路５に入
力した場合のノイズ開始時と終了時の出力値を記録し、
この出力値の平均値を求める。この平均値が閾値決定の
指標に相当する。この算出値に対して、“ノイズのない
対象データ”のノイズ換算値を求めるために、このデー
タをノイズキヤンセル回路５に入力し、ＣＩＣフイルタ
３３の出力の最大値を求める。そしてこの最大値を先に
求めた“ノイズデータ”算出値（出力特性）により逆換
算し、量子化回路３４の閾値ｘを求める。In this embodiment, "Gaussian white noise" which is uniform with respect to frequency is used as "noise data", a data string in which this noise bursts and appears is created, and this data string is generated by the noise cancel circuit 5. Record the output value at the start and end of noise when input to
The average value of these output values is calculated. This average value corresponds to the threshold determination index. In order to obtain a noise conversion value of "target data having no noise" with respect to this calculated value, this data is input to the noise cancel circuit 5, and the maximum value of the output of the CIC filter 33 is obtained. Then, this maximum value is inversely converted by the previously calculated "noise data" calculated value (output characteristic) to obtain the threshold value x of the quantization circuit 34.
【００２０】なお、本実施例の如くに“ノイズの無い対
象データ”が心電図データの様に個人差を持つデータで
ある場合には、予め個人データのノイズ換算値を求めて
おくことにより、より信頼性の高い結果が得られる。ま
た、前もつてノイズ換算値が得られない様な場合には、
一般的な使用に耐えるようにするために多数の対象デー
タを統計的に処理し、ノイズ換算値を得ておく必要があ
る。この場合には、対象データの分散により、ノイズキ
ヤンセル能力の限界が決定されることになる。When the "noise-free target data" is data having individual differences such as electrocardiogram data as in the present embodiment, it is possible to obtain a noise conversion value of individual data in advance. Reliable results are obtained. If the noise conversion value cannot be obtained even before,
In order to endure general use, it is necessary to statistically process a large number of target data and obtain a noise conversion value. In this case, the variance of the target data determines the limit of the noise canceling ability.
【００２１】図２に示す微分回路３１およびＣＩＣフイ
ルタ３３の詳細構成を図３に示す。本実施例において
は、サンプリング周波数１８０ヘルツの心電図データの
うち有効成分を９０ヘルツまでとみなし、微分回路３１
を合計４段の線形フイルタ回路で構成して一種のＨＰＦ
構成とし、高周波成分を増幅し、ノイズ成分の平均値パ
ワーを得るようにしている。このため、各段を、入力信
号をＺ-1して加算器の“−”入力に、入力信号を加算器
の“＋”入力に接続して線形フイルタを構成し、加算器
出力にデジタルシステムでは１ビツトシフトに相当する
（１／２）増幅器を備えて入力信号との対応をとるとい
う、加算器のみで乗算器を含まない演算回路構成として
いる。本実施例では、種々の実験を行なつた結果、必要
最小限の微分回路の段数として４段を採用した。The detailed construction of the differentiating circuit 31 and the CIC filter 33 shown in FIG. 2 is shown in FIG. In the present embodiment, the effective component of the electrocardiogram data having a sampling frequency of 180 Hertz is regarded as 90 Hertz, and the differentiating circuit 31 is used.
Is composed of a total of 4 stages of linear filter circuits
The configuration is such that the high frequency component is amplified and the average value power of the noise component is obtained. Therefore, each stage is connected to the "-" input of the adder by Z -1 of the input signal and connected to the "+" input of the adder to form a linear filter, and a digital system is provided at the output of the adder. In the above, an arithmetic circuit configuration is provided in which a (1/2) amplifier corresponding to one bit shift is provided to correspond to an input signal, that is, an adder only and a multiplier is not included. In this embodiment, as a result of various experiments, four stages were adopted as the minimum required number of stages of differentiating circuits.
【００２２】また、ＣＩＣフイルタ３３は、加算器の
“−”入力には原入力信号をＺ-128した信号を入力し、
“＋”入力に原入力信号および加算器の出力信号をＺ-1
した信号を入力した構成とし、加算器出力にデジタルシ
ステムでは７ビツトシフトに相当する（１／１２８）増
幅器を接続して原入力信号との対応を取つている。以上
説明したように、このノイズ検出回路及びノイズ除去フ
イルタを、加算器を基本とした乗算器を必要としない構
成とすることにより、高速処理が実現し、回路構成も簡
単なものと出来る。しかも、必要な場合のみノイズ除去
フイルタ２０を通過させることにより、不必要に入力信
号が変形するようなこともなく、忠実度の高い圧縮処理
が可能となる。なお、図３においては、Ｚの個数も実験
結果に基づいて決定されており、それぞれの必要最小限
の構成となつている。The CIC filter 33 also inputs a signal obtained by Z -128 the original input signal to the "-" input of the adder,
The original input signal and the adder output signal are input to the "+" input by Z -1.
In the digital system, a (1/128) amplifier corresponding to 7 bits shift is connected to the output of the adder to correspond to the original input signal. As described above, by arranging the noise detecting circuit and the noise removing filter so as not to require a multiplier based on an adder, high speed processing can be realized and the circuit structure can be simplified. Moreover, by passing the noise removal filter 20 only when necessary, high-fidelity compression processing can be performed without unnecessarily deforming the input signal. In addition, in FIG. 3, the number of Zs is also determined based on the experimental results, and each has the necessary minimum configuration.
【００２３】本実施例のホルタ心電計においては、符号
化圧縮処理を行う前に不要なノイズ成分を有効に除去す
ることにより、圧縮率を上げることができる。このた
め、好適な圧縮処理を施すことができ、小容量のＲＡＭ
に長時間の心電図情報を記憶させることができる。この
場合においても、心電図情報をデジタルデータの形で記
憶するため、解像度が劣化したり、波形歪が起こること
もない、ホルタ心電計とすることができる。しかも、可
動部もないため、振動などにも強く、小型軽量化も容易
であり、価格も低く抑えることが可能となる。In the Holter electrocardiograph of the present embodiment, the compression rate can be increased by effectively removing the unnecessary noise component before performing the coding compression processing. Therefore, a suitable compression process can be performed, and a small capacity RAM
It is possible to store the electrocardiogram information for a long time. Even in this case, since the electrocardiogram information is stored in the form of digital data, it is possible to provide a Holter electrocardiograph which does not cause deterioration in resolution and waveform distortion. Moreover, since there are no moving parts, it is resistant to vibration and the like, and it is easy to reduce the size and weight, and the price can be kept low.
【００２４】[0024]
【発明の効果】以上説明した様に本発明によれば、心電
図情報中のノイズを有効にキヤンセルすることができ、
その後に圧縮処理を行なう際に十分な圧縮を行なうこと
ができ、長時間の心電図情報を記録する際に必要とする
記憶容量を低く抑えることが可能となる。As described above, according to the present invention, noise in electrocardiogram information can be effectively canceled.
Sufficient compression can be performed when the compression process is performed thereafter, and the storage capacity required for recording the electrocardiogram information for a long time can be suppressed to a low level.
【００２５】しかも、ノイズの検出を加算器を基本とし
た乗算器を必要としない構成で行なうことにより、高速
処理が実現し、回路構成も簡単なものと出来る。Moreover, high-speed processing can be realized and the circuit structure can be simplified by detecting noise with a structure that does not require a multiplier based on an adder.
【図１】本発明に係る一実施例のホルタ心電計のブロツ
ク図である。FIG. 1 is a block diagram of a Holter electrocardiograph according to an embodiment of the present invention.
【図２】図１に示すノイズキヤンセル回路の詳細構成を
示すブロツク図である。FIG. 2 is a block diagram showing a detailed configuration of a noise cancel circuit shown in FIG.
【図３】図２の詳細構成を示す図である。FIG. 3 is a diagram showing a detailed configuration of FIG.
【図４】本実施例のノイズキヤンセル回路の入出力波形
を示す図である。FIG. 4 is a diagram showing input / output waveforms of the noise cancel circuit of the present embodiment.
１〜３ 生体誘導電極（ピツクアツプ） ４ アンプ回路 ５ ノイズキヤンセル回路 ６ 符号化圧縮回路 ７ 書き込み回路 ８ ＲＡＭ １０ 位相調整回路 ２０ ノイズ除去フイルタ回路 ３０ ノイズ検出回路 ３１ 微分回路 ３２ 絶対値回路 ３３ ＣＩＣ回路 ３４ 量子化回路 ４０ 切替回路 ５０ 基線変動除去フイルタ回路 1 to 3 biological induction electrode (pickup) 4 amplifier circuit 5 noise cancel circuit 6 coding compression circuit 7 writing circuit 8 RAM 10 phase adjusting circuit 20 noise removal filter circuit 30 noise detection circuit 31 differential circuit 32 absolute value circuit 33 CIC circuit 34 Quantization circuit 40 Switching circuit 50 Baseline fluctuation removal filter circuit
Claims (3)
イズ成分のパワーを検出する検出手段と、 該検出手段の検出パワーが所定以上の場合には前記誘導
電極よりの心電図情報に対してノイズ除去フイルタを介
してノイズキヤンセルを行なつた後に基線変動除去を行
ない、前記検出手段よりの検出パワーが所定以上でない
場合にはノイズ除去フイルタを介さずそのまま基線変動
除去を行なうノイズ除去手段とを備えることを特徴とす
る心電計におけるノイズキヤンセル回路。1. A detection means for detecting the power of a noise component contained in electrocardiogram information from the induction electrode, and noise removal for the electrocardiogram information from the induction electrode when the detection power of the detection means is above a predetermined level. And a noise removing means for removing the baseline fluctuation after performing the noise cancel via the filter, and for removing the baseline fluctuation without passing through the noise removing filter when the detection power from the detecting means is not more than a predetermined value. Noise cancel circuit in electrocardiograph characterized by.
る偶数段の微分回路と、該微分回路よりの微分波形の絶
対値を出力する絶対値回路と、該絶対値回路に接続され
たＣＩＣフイルタ回路で構成することを特徴とする請求
項１記載の心電計におけるノイズキヤンセル回路。2. The detection means includes an even number stage differentiating circuit having at least two stages, an absolute value circuit for outputting an absolute value of a differential waveform from the differentiating circuit, and a CIC filter connected to the absolute value circuit. The noise cancel circuit in the electrocardiograph according to claim 1, wherein the noise cancel circuit is configured by a circuit.
イズ成分のパワーを検出し、検出したノイズ成分パワー
が所定以上の場合には前記誘導電極よりの心電図情報に
対してノイズ除去フイルタを介してノイズキヤンセルを
行なつた後に基線変動除去を行ない、前記検出したノイ
ズ成分パワーが所定以上でない場合にはノイズ除去フイ
ルタを介さずそのまま基線変動除去することを特徴とす
る心電計におけるノイズキヤンセル方法。3. The power of the noise component contained in the electrocardiogram information from the induction electrode is detected, and when the detected noise component power is equal to or higher than a predetermined level, the electrocardiogram information from the induction electrode is passed through a noise removal filter. A noise cancel method in an electrocardiograph, characterized in that the baseline fluctuation is removed after performing the noise cancellation, and if the detected noise component power is not more than a predetermined value, the baseline fluctuation is removed as it is without passing through the noise cancellation filter.
Priority Applications (1)
| Application Number | Priority Date | Filing Date | Title |
|---|---|---|---|
| JP5068009A JPH06277189A (en) | 1993-03-26 | 1993-03-26 | Method for cancelling noise and noise cancelling circuit in electrocardiograph |
Applications Claiming Priority (1)
| Application Number | Priority Date | Filing Date | Title |
|---|---|---|---|
| JP5068009A JPH06277189A (en) | 1993-03-26 | 1993-03-26 | Method for cancelling noise and noise cancelling circuit in electrocardiograph |
Publications (1)
| Publication Number | Publication Date |
|---|---|
| JPH06277189A true JPH06277189A (en) | 1994-10-04 |
Family
ID=13361429
Family Applications (1)
| Application Number | Title | Priority Date | Filing Date |
|---|---|---|---|
| JP5068009A Pending JPH06277189A (en) | 1993-03-26 | 1993-03-26 | Method for cancelling noise and noise cancelling circuit in electrocardiograph |
Country Status (1)
| Country | Link |
|---|---|
| JP (1) | JPH06277189A (en) |
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|---|---|---|---|---|
| JP2006512168A (en) * | 2002-12-30 | 2006-04-13 | カーディアック・ペースメーカーズ・インコーポレーテッド | Method and apparatus for monitoring diastolic hemodynamics |
| US7844334B2 (en) | 2003-11-06 | 2010-11-30 | Cardiac Pacemakers, Inc. | Dual-use sensor for rate responsive pacing and heart sound monitoring |
| US8103335B2 (en) | 2004-12-22 | 2012-01-24 | Nihon Kohden Corporation | Cardiogram waveform correcting and displaying device and a method of correcting and displaying cardiogram waveforms |
| JP2013059439A (en) * | 2011-09-13 | 2013-04-04 | Nippon Koden Corp | Biological signal measurement device |
| US8951205B2 (en) | 2002-12-30 | 2015-02-10 | Cardiac Pacemakers, Inc. | Method and apparatus for detecting atrial filling pressure |
| CN105764413A (en) * | 2013-09-30 | 2016-07-13 | 赛佛欧普手术有限公司 | Systems and methods for preventing contamination of recorded biological signals during surgery |
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| US11064949B2 (en) | 2016-05-26 | 2021-07-20 | Samsung Electronics Co., Ltd. | Method and apparatus to remove noise from electrocardiography (ECG) sensor signal |
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|---|---|---|---|---|
| JP2006512168A (en) * | 2002-12-30 | 2006-04-13 | カーディアック・ペースメーカーズ・インコーポレーテッド | Method and apparatus for monitoring diastolic hemodynamics |
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| US8403860B2 (en) | 2002-12-30 | 2013-03-26 | Cardiac Pacemakers, Inc. | Method and apparatus for detecting atrial filling pressure |
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| US7844334B2 (en) | 2003-11-06 | 2010-11-30 | Cardiac Pacemakers, Inc. | Dual-use sensor for rate responsive pacing and heart sound monitoring |
| US8103335B2 (en) | 2004-12-22 | 2012-01-24 | Nihon Kohden Corporation | Cardiogram waveform correcting and displaying device and a method of correcting and displaying cardiogram waveforms |
| JP2013059439A (en) * | 2011-09-13 | 2013-04-04 | Nippon Koden Corp | Biological signal measurement device |
| JP2016537038A (en) * | 2013-09-30 | 2016-12-01 | セーフオプ サージカル インコーポレイテッド | System and method for preventing contamination of biological signals recorded during surgery |
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| CN113080875A (en) * | 2013-09-30 | 2021-07-09 | 赛佛欧普手术有限公司 | System and method for preventing recorded bio-signal interference during surgery |
| US11406303B2 (en) | 2013-09-30 | 2022-08-09 | Safeop Surgical, Inc. | Systems and methods for preventing contamination of recorded biological signals during surgery |
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