JP3722404B2 - High frequency ventilator - Google Patents

Description

（n：ガスの粘性係数，r：半径，I：長さ）
【００９０】
第1分岐抵抗Ｒ１＋…＋第10分岐抵抗Ｒ１０については、次のように考える。図９に示した線図によれば、第1〜第10分岐までの各分岐の気道累積面積（同次元の全ての気道の断面積の合計）は、ほぼ等しい。患者によって様々であるが、ヒトの第1〜第4分岐までの気道累積面積の代表例を図８の値を参考とする。第1〜第4分岐までの気道累積断面積は、平均で46[cm²]程度とできる。これが第1〜第10分岐まで、ほぼ等しい累積面積でいくとすれば、第1〜第10分岐までは、直径3.8[cm]の一本の管と同等の空気抵抗を持つと仮定してもよい。そして、第1〜第10分岐の長さは約20[cm]程度と仮定すると、肺内の第1分岐〜第10分岐までの気道抵抗の総和は次式（５）で表すことができる。
【００９１】
第1分岐抵抗Ｒ１＋…＋第10分岐抵抗Ｒ１０
＝（8×ｎ×ｌ）/πｒ⁴＝（8×ｎ×20）/(π・1.4⁴)＝41.6ｎ/π＝13.3ｎ …（５）
【００９２】
ガスの粘性係数ｎは一定であるから、「気管内挿入管８１抵抗Ｒｃ」は、「第1分岐抵抗Ｒ１＋…＋第10分岐抵抗Ｒ１０」に対して、2984n/13.3ｎ＝約224倍となる。
【００９３】
よって、気管内挿入管８１抵抗Ｒｃ＞＞第1分岐抵抗Ｒ１＋…＋第10分岐抵抗Ｒ１０となるので、気道抵抗Ｒは気管内挿入管８１抵抗Ｒｃでほとんど生じ、ほぼ同等と考えてもよい。特に、高頻度人工呼吸のように速くて小さな流動を生じる呼吸方法は、空気の流れが深い分岐まで直接至らないため、実際に空気の流れが起っているのは、主に次数の少ない分枝までである。そのため空気の流れがあるときに考慮すべきＲの項は、「気管内挿入管８１抵抗Ｒｃ＞＞第1分岐以降の総分岐抵抗」の傾向がさらに強くなる。なお、高頻度人工呼吸法では、直接的な空気の流れにて酸素を行き渡らせるのではなく酸素の拡散により換気を行うので、深い分枝まで直接的な流れが到達しない場合でも、換気は充分に行われる。
【００９４】
このように、高頻度人工呼吸法の換気法を採用している上では、気道抵抗Ｒは、使用する「気管内挿入管８１」の空気抵抗によって代表させて問題ない。
【００９５】
これらの理由により（１）式のΔＰ＝Ｒ・ΔＶ/ｄｔ＋ΔＶ/ＣのＲ項を「気管内挿入管８１」の空気抵抗である気道抵抗Ｒｃとし、Ｃ項を約200[ml/cmH20]の硬質タンクの「モデル肺８２」のコンプライアンスＣとすることで、単純モデル化できる。
【００９６】
これらモデル化したHFO回路を用いて、「モデル肺８２」に発生している圧力Ｐ１の変化と同等もしくは相似の変化（前述した一時間数に基づく変化）をする「圧力検出容器８５」と「検出用分岐管８４」とを製作すれば、同じ回路構成のHFO呼吸器を患者に用いた場合、間接的に「実際の肺」の圧力Ｐ３の変化が、「圧力検出容器８５」の圧力Ｐ２の変化を計測することで知ることができる。
【００９７】
上述のことを式でまとめると次のようになる。HFO時の患者の肺内圧変化をΔＰとし、肺内に出入りしたガスの変化量（流量）をΔＶとし、HFO時の肺の気道抵抗をＲ、肺のコンプライアンスをＣとすると前述した式（１）が成立する。
【００９８】
このとき、高頻度人工呼吸時では、前述したように、Ｒの項は患者に使用している「気管内挿入管８１」の空気抵抗Ｒｃと近似的に置き換えてもほぼ問題ない。さらに、コンプライアンスＣはモデル肺８２のコンプライアンス（=200[ml/cmH2O]）とすることができるので次式（６）が成立する。
【００９９】
ΔＰ＝Ｒｃ・（ｄΔＶ/ｄｔ）＋ΔＶ/200 …（６）
【０１００】
〔圧力検出容器と検出用分岐管〕
ここで、Ｒｄの空気抵抗を持つ「検出用分岐管８４」とＣ’のコンプライアンスを持つ「圧力検出容器８５」を、図１に示す「モデル肺８２と圧力検出容器８５の関係」のように同じHFO回路内で使用する。但し、ＲｄはＲｃと、Ｃ’はＣと、後述する相対的な関係にあるものとする。「圧力検出容器８５」で起るボリューム変化をΔＶ’とすると次式（７）が成り立つ。
【０１０１】
ΔＰ＝Ｒｄ・（ｄΔＶ’/ｄｔ）＋ΔＶ’/Ｃ’ …（７）
【０１０２】
「モデル肺８２」で起る圧変化Ｐ１は、前述した式（６）による圧変化をし、「圧力検出容器８５」で起る圧変化Ｐ２は上式（７）による圧変化をする。上式（７）の全ての項は前式（６）と相対的な比例関係にあるため、Ｐ１とＰ２の関係は、次式になる。
【０１０３】
Ｐ１＝ｋ×Ｐ２＋ｍ …（８）
（k，mは定数）
【０１０４】
ここで、HFO時の実際の患者の肺の内圧変化をＰalvとすると、HFO特有の換気メカニズムから、実際の肺は「モデル肺８２」として単純モデル化できるため、患者の肺の内圧変化Ｐalvは次式（９）で表すことができる。
【０１０５】
Ｐalv≒Ｐ１ …（９）
【０１０６】
そこで、実際の患者がつながっている高頻度人工呼吸器１２が図２の状態（人工呼吸モード）で「圧力検出容器８５」の圧力変化Ｐ２を求めれば、式（８）、（９）により患者の肺の内圧変化Ｐalvが求められる。
【０１０７】
Ｐalv≒Ｐ１＝ｋ×Ｐ２＋ｍ …（１０）
【０１０８】
上式（１０）で患者の肺内圧変化が間接的に計測できる。ここで、圧力検出容器８５及び検出用分岐管８４の設置位置は、できるだけモデル肺８２及び気管内挿入管８１と近接することが望ましいが、他の位置でも良い。
【０１０９】
以上は、正常な肺モデルの静的コンプライアンスＣ＝約200[ml/cmH20]で「モデル肺８２」を代表値として考えたが、病的な肺のコンプライアンスは、正常肺の静的コンプライアンスＣに係数ｎをかければ表現できる。病的肺のコンプライアンスをＣ''とすれば次式（１１）で表すことができる。
【０１１０】
Ｃ''＝ｎＣ …（１１）
（ｎは定数）
【０１１１】
よって、病的肺の内圧変化は、式（６）を変形して次式（１２）となる。
【０１１２】
ΔＰ＝Ｒｃ・（ｄΔＶ/ｄｔ）＋ΔＶ/200ｎ …（１２）
【０１１３】
式（１２）の第2項は、Ｃの比例倍の関係にある項であるため、（１０）式は次のようになる。
【０１１４】
Ｐalv≒Ｐ１＝ｋ'×Ｐ２＋ｍ …（１３）
（ｋ'は係数ｎを考慮したｋの比例倍定数）
【０１１５】
ｋとｋ'の関係は、病的肺の静的コンプライアンスＣ''を持つ「モデル肺８２’」と正常な肺のコンプライアンスを持つ「モデル肺８２」とを用いて比較すれば、比例関係式が導かれる。よって、実際の臨床においては患者の診療時に、病的肺の静的コンプライアンスＣ''をあらかじめ計測し、その値と正常肺との関係を（１１）式により求めれば、どのｋ'を使用すればよいかわかる。このｋ'を用いれば、病的肺においても、Ｐ２から患者の肺内圧変化Ｐ１が（１３）式より求まる。
【０１１６】
〔一回換気量の算出〕
さらに、次のような計算式から、実際の患者の肺にどれだけの一回換気が出入りしているかが推定できる。
【０１１７】
基準として、「モデル肺８２」に定量のガス100[ml]を瞬間的に押し込んだ時の圧力変化をΔＰa（定数）とする（ΔＰａ＝100/Ｃ、例えばＣ＝200[ml/cmH2O]であればΔＰａ＝0.5[cmH2O]）。高頻度人工呼吸法の換気により「モデル肺８２」にΔＰ１の圧振幅が発生すれば、ボイルの法則を用いて、次式（１４）の関係が成立する。
【０１１８】
ΔＰa/100＝ΔＰ１/ΔＶ＝１/Ｃ …（１４）
【０１１９】
また、高頻度人工呼吸時の圧力検出容器のＰ２の圧振幅変化ΔＰ２はΔＰ１と（１０）式の関係があるので、次式（１５）が成立する。
【０１２０】
ΔＰa/100＝ΔＰ１/ΔＶ＝(ｋ・ΔＰ２＋ｍ)/ΔＶ＝１/Ｃ …（１５）
【０１２１】
よって、「モデル肺８２」のHFO時のボリューム変化であるΔＶ（換気量）が求まる。
【０１２２】
ΔＶ＝100×(ｋ・ΔＰ２＋ｍ)/ΔＰa＝Ｃ・(ｋ・ΔＰ２＋ｍ) …（１６）
【０１２３】
「モデル肺８２」と「実際の肺」は、流体力学的にほぼ同じであることから、実際の患者の肺の換気量もほぼΔＶとすることができる。つまり、準備作業段階の測定で係数ｋ，ｍを求め、高頻度人工呼吸時に圧変化ΔＰ２を計測すれば、患者の肺の換気量も式（１６）を用いて、ほぼ推定できる。
【０１２４】
なお、コンプライアンス値Ｃは前述した▲１▼〜▲３▼の理由により、高頻度人工呼吸時における圧力検出に際しては、気道抵抗Ｒと比較してモデル肺８２及び圧力検出容器８５の圧力検出に対する影響は小さく、これらの内部圧力の変化の主な特定要素は、気管内挿入管８１及び検出用分岐管８４の気道抵抗に左右されているといえる。従って、モデル肺８２の容積の設定に際しては、コンプライアンス値Ｃに基づいて算出した最も好適な容積から若干の誤差(±150[％])があっても検出圧力に大きな誤差は生じない（より精度を高くするためには最も好適な値に設定することが当然望ましい）。但し、式（１６）から換気量ΔＶを算出するときにはコンプライアンスＣは正確でなくてはならない。
【０１２５】
【第１実施例】
式（８）の関係を、前述した高頻度人工呼吸器１２のシステムを利用して測定した結果を説明する（図２参照）。
【０１２６】
まず、吸気導入部６２にて高純度の酸素（約100％）と空気を混合して医者の設定した値の高濃度酸素の混合ガスを作り出す。同時に流量を医者が設定した値（10〜60[l/min]）の吸気（フレッシュガス）を患者回路内に定常流で流す。この定常流に振動空気圧付勢部５０から振動を加え、空気の振動波を作り出す。流れを伴った空気振動波が、患者Ｘの肺のガス交換を行なう。
【０１２７】
また、排出経路７０の出口開放面積を流量調節バルブ６０７で可変制御させることにより、患者の肺に負担をかけないように医者が設定した適切な平均圧力（通常、平均気道内圧という。平均気道内圧Ｐmean＝5〜15[cmH20]）を回路内で保ち、患者の肺にも一定の陽圧を付加しながら換気をしている。この準備作業段階での代表的な「モデル肺８２」の圧力波形例を図１１に示す。
【０１２８】
さらに、「モデル肺８２」のＰ１と「圧力検出容器８５」のＰ２の圧力変化を比較した実際の実験結果を説明する。実験条件は、成人患者を想定した時の「モデル肺８２」及び「気管内挿入管８１」を使う。「モデル肺８２」及び「気管内挿入管８１」と相対的な気道抵抗とコンプライアンスをもつ「圧力検出容器８５」及び「検出用分岐管８４」を試作する。
【０１２９】
実験条件は以下の如く設定した。
「モデル肺８２」………………40[l]タンク
「気管内挿入管８１」…………内径8[mm]のチューブ８１，（長さ30[cm]）
「圧力検出容器８５」…………10[ml]容器
「検出用分岐管８４」…………内径0.5[mm]のチューブ８１，（長さ30[cm]）
上記実験条件のもとに、振動空気圧の周波数を5[Hz]と15[Hz]とに設定して実験を行った。5[Hz]の圧力波形データを図１２に示し、実験結果を図１３に示す。また、15[Hz]の圧力波形データを図１４に示し、実験結果を図１５に示す。
【０１３０】
Ｐ１とＰ２の関係は、いずれの周期の場合でも、相対的な関係で安定しており、前式（８）で示すことが可能である。
【０１３１】
【第２実施例】
また、医者が、HFO換気を行なっている患者の肺を計測するために、対応した「モデル肺８２＋気管内挿入管８１」をどのように選定するかが問題になる。
【０１３２】
先に延べたように、患者に対して使用する「気管内挿入管８１」は、患者の年齢により気管内挿入管８１のサイズ（内径）と挿管深さ（チューブの長さ）が決まる（図４参照）。これは、人工呼吸管理をする際に利用する、医学的な統計データに基づいた指標値として、一般的に扱われている。
【０１３３】
また、患者の年齢が決まれば、肺のコンプライアンスも医学的な統計データから、ほぼ値が求められている。肺の病状により病的肺のコンプライアンスも医学的な統計データから、ほぼ値が求められている。また、臨床的に計測できる装置もあるので、簡単に求められる。
【０１３４】
ここで、実験条件を小児患者（3歳程度）を想定した時の「モデル肺８２’」及び「気管内挿入管８１’」にする。「モデル肺８２'」及び「気管内挿入管８１’」と相対的な気道抵抗とコンプライアンスをもつ「圧力検出容器８５'」及び「検出用分岐管８４'」を試作する。
【０１３５】
実験条件は以下の如く設定した。
「モデル肺８２'」……………20[l]タンク
「気管内挿入管８１'」………内径4[mm]気管内挿入管８１，（長さ19[cm]）
「圧力検出容器８５'」………12[ml]容器
「検出用分岐管８４'」………内径0.5[mm]気管内挿入管８１，（長さ120[cm]）
上記実験条件のもとに、振動空気圧の周波数を5[Hz]と15[Hz]とに設定して実験を行った。5[Hz]の実験結果を図１６に示す。また、15[Hz]の実験結果を図１７に示す。
【０１３６】
Ｐ１とＰ２の関係は、いずれの周期の場合でも、相対的な関係で安定しており、前式（８）と同様の式（８）’で示すことが可能である。
【０１３７】
Ｐ１’＝ｋ’×Ｐ２’＋ｍ’ …（８）’
（k’，m’は定数）
【０１３８】
以上のように実際の実験結果において、どのような対象患者の肺に対しても、患者の肺とコンプライアンス値の等しいモデル肺とこれと相似関係にある圧力検出容器とから、患者の肺内圧変化を相対的に知ることができる。
【０１３９】
また、患者の病状に合わせてP2に乗算する係数を考慮すれば正確に肺内圧が計測できることも実験結果よりわかる。
【０１４０】
よって、「気管内挿入管８１」の各部の寸法と「患者の肺コンプライアンス」とは、いずれも患者の年齢で任意に決まるので、これらの値を年齢別に組み合わせたセットのデータとして取り扱っても良い。また、それに相対的に合わせた「検出用分岐管８４」の長さ及び内径の寸法と「圧力検出容器８５のコンプライアンス（或いは圧力検出容器８５の容積）」の値も年齢別に組み合わせたセットのデータとして取り扱っても良い。
【０１４１】
これにより、医者は、HFOを施す患者に使用する気管内挿入管８１を決めさえすれば、その肺内圧を測定する「圧力検出容器８５」と「検出用分岐管８４」のセットも、医者が迷うことなく同時に決まってしまうため、選定が楽であり、扱いが容易となる。また、患者の年齢に合わせた、「圧力検出容器８５」と「検出用分岐管８４」があることで、より正確に患者の肺内圧を間接的に、かつ正確に計測できることになる。
【０１４２】
また、式（８）の係数ｋがほぼ1に等しくなるように圧力検出容器８５及び検出用分岐管８４を試作すれば、ｍの項による単なる圧力値のオフセットしか生じないので、これにより圧波形のカーブ変化がモデル肺用圧力センサ出力と対応圧力検出センサ出力とで同じになるため、医者が観察し易く、肺内圧力及び換気量の計算も容易となる。図１２、図１４の波形データは、係数ｋをほぼ1になるように圧力検出容器８５及び検出用分岐管８４を設計したため、波形データがほぼ同じである。
【０１４３】
またこの場合の波形の「ピ一クtoピ一ク」は、「モデル肺８２」と「圧力検出容器８５」で一致するため、前式（１６）から、間接的に患者の換気量が簡単に算出でき、高頻度人工呼吸法の人工呼吸の管理指標となり、医者としては患者の治療管理が容易になる。
【０１４４】
【発明の効果】
本願発明では、検出用分岐管の内径及び長さ並びに圧力検出容器の容積を、高頻度人工呼吸時の気管内挿入管の挿入側先端部圧力と圧力検出容器の内部圧力とが一次関数の関係を生ぜしめる値に設定されている。そして、コントローラでは、患者に高頻度人工呼吸を行っている際に、圧力検出容器の内部の検出圧力から患者の肺内圧力を算出する。
【０１４５】
従って、本願発明により、従来には不可能であった、高頻度人工呼吸の最中における患者の肺内の圧力変化及び換気量の測定が可能となり、熟練と経験を要することなく換気状態の観測が可能となり、また、従来の如く目視に頼る必要がないので、より正確な状態を把握することが可能となり、常により好適な高頻度人工呼吸の維持が可能となった。
【０１４６】
また、上述の一次関数を入力する操作盤を設けることにより、選定された検出用分岐管及び圧力検出容器に応じて決定される一次関数を自在に入力することが可能となる。
【０１４７】
さらに、算出された肺内圧力を表示する表示部を設けることにより、高頻度人工呼吸時における肺内圧力を容易且つ迅速に知ることが可能となる。
【０１４８】
また同様にして、換気量算出部を備えることにより、外部処理装置を不要とし、本願発明の構成のみにより、換気量の把握が可能となる。
【０１４９】
本発明は以上のように構成され機能するので、これによると、従来にない優れた高頻度人工呼吸器を提供することができる。
【図面の簡単な説明】
【図１】本実施形態たる高頻度人工呼吸器の構成を示すブロック図であって、人工呼吸時の状態を示す。
【図２】本実施形態たる高頻度人工呼吸器の構成を示すブロック図であって、人工呼吸を行う前の準備作業段階の状態を示す。
【図３】図１，２で開示した流量調節バルブの詳細を示す断面図である。
【図４】年齢に応じた気管内挿入管の寸法例を示す図表である。
【図５】高頻度人工呼吸器の制御系を示すブロック図である。
【図６】吸気供給時と呼気排出時とにおける「内圧−換気量」の関係を示す線図である。
【図７】人間の肺の模型を示す説明図である。
【図８】「気道の分岐構成」を示す説明図である。
【図９】分岐の次元と各次元ごとの気道断面積の総和との関係を示す線図である。
【図１０】分岐の次元と各次元ごとの気管による気道抵抗の関係を示す説明図である。
【図１１】本実施形態による高頻度人工呼吸を行った際の代表的なモデル肺の内部圧力の波形例を示す線図である。
【図１２】図１に示す高頻度人工呼吸器にて振動空気圧を周波数5Hzとしたときの第一及び対応圧力検出センサの検出圧力を示す線図である。
【図１３】図１に示す高頻度人工呼吸器にて振動空気圧を周波数5Hzとしたときの試験結果を示す図表である。
【図１４】図１に示す高頻度人工呼吸器にて振動空気圧を周波数15Hzとしたときの第一及び対応圧力検出センサの検出圧力を示す線図である。
【図１５】図１に示す高頻度人工呼吸器にて振動空気圧を周波数15Hzとしたときの試験結果を示す図表である。
【図１６】図１に示す高頻度人工呼吸器の各チューブ及び各密閉容器を小児を対象とする設定とした場合であって、振動空気圧を周波数5Hzとしたときの試験結果を示す図表である。
【図１７】図１に示す高頻度人工呼吸器の各チューブ及び各密閉容器を小児を対象とする設定とした場合であって、振動空気圧を周波数15Hzとしたときの試験結果を示す図表である。
【図１８】従来の高頻度人工呼吸器の構成を示すブロック図である。
【符号の説明】
１２ 高頻度人工呼吸器
４０ コントローラ
４２ 記憶部
４３ 操作盤
４４ 表示部
４７ 肺内圧力算出部
４８ 換気量算出部
５０ 振動空気圧付勢部
６０ 患者側経路
６２ 吸気導入部
７０ 排出経路
８１ 気管内挿入管
８２ モデル肺
８３ モデル肺用圧力センサ
８４ 検出用分岐管
８５ 圧力検出容器
８６ 対応圧力検出センサ
Ｘ 患者[0001]
BACKGROUND OF THE INVENTION
The present invention relates to a high-frequency ventilator, and more particularly to a high-frequency ventilator that can supply oxygen while observing a patient's intrapulmonary pressure.
[0002]
[Prior art]
As shown in FIG. 18, the conventional high-frequency ventilator 200 contains high-concentration oxygen flowing through a fluid circuit system that branches from the oxygen supply source 201 to the patient X side and the exhaust side via the three-way branch pipe 202. The oscillating air pressure unit 203 urges the oscillating air pressure at a high frequency (about 3 to 15 [Hz]) with respect to the intake air (normal flow rate 10 to 30 [l / min], maximum 60 [l / min]). Then, oxygen is supplied into the lungs of patient X. At this time, the average pressure applied to the lungs of the patient X is controlled by the closed area of the rubber valve of the exhalation valve 204 provided at the exhalation outlet, and the normal average pressure is maintained at 5 to 15 [cmH2O]. Set to.
[0003]
The oxygen supply principle of the high frequency ventilator 200 will be described. First, when high-frequency oscillating air pressure is applied to the inspiratory air supplied to the patient, the volume of the inhaled air containing the carbon dioxide in the patient's lungs (hereinafter referred to as exhalation) is reduced by the pressure amplitude of the inspiratory air. Ventilation (convective gas exchange) occurs, and due to the effect of diffusion movement by the vibration of inspiration, inhalation enters the lung and exhalation in the lung is led out of the lung (patient's mouth). The subsequent inspiration performs the above-described ventilation and also has an action of sending the exhaled air derived from the lungs to the exhaust port side. This makes it possible to always maintain a constant oxygen concentration in the patient's lungs.
[0004]
[Problems to be solved by the invention]
In a high-frequency ventilator, gas ventilation that switches in a fast cycle of 3 to 15 [Hz] is performed on the patient's lungs in both directions of inspiration and exhaust. It was difficult to measure the flow (or volume) with a normal flow sensor. Therefore, it has been difficult to detect the amount of ventilation in the patient's lungs.
[0005]
For example, a hot wire type sensor can be cited as a flow rate sensor with good responsiveness. Focusing on the fact that the hot wire heated by energization is cooled in proportion to the flow rate of gas passing through the hot wire, the hot wire sensor detects the flow rate. Since there is a proportional relationship, the flow rate is detected from the current value.
[0006]
However, regardless of the direction of gas flow, this method can measure the total amount of gas that passes because the heat ray is cooled if there is a large flow rate of gas that passes, but it can measure which direction the gas is flowing. Can not. For this reason, a hot-wire sensor can measure the instantaneous minute flow rate (flow rate per unit time) at the sensor position for gas flow in both directions of inspiration and exhalation, but it can actually measure in the lungs. On the other hand, it was not possible to know how much inspiration was flowing in or how much exhalation was being exhausted.
[0007]
There is also a method of measuring the gas amount by measuring the differential pressure across the orifice provided in the middle of the pipe through which the gas flows. With this method, if the reference pressure is fixed to any pressure before or after the orifice, the front-rear differential pressure ΔP can be obtained as a plus or minus depending on the difference in the gas flow direction, so that the gas flow direction can be determined. .
[0008]
However, in this method, an orifice that generates a differential pressure in the middle of the pipe becomes a ventilation resistance, which hinders inhalation supply or exhalation. In addition, if the orifice diameter is set to be large in order to suppress the ventilation resistance, the differential pressure before and after becomes a few [cmH2O], and since there are few pressure sensors that detect such a fine pressure with high accuracy, an accurate ΔP can be obtained as a result. There was no inconvenience.
[0009]
Further, if it is possible to detect a change in pressure in the lungs of the patient during artificial respiration, the amount of gas exchange can be calculated. However, with the above-mentioned high-frequency ventilator, the pressure can be detected only at the patient's mouth, and the patient's intrapulmonary pressure does not match the mouth pressure. I couldn't.
[0010]
For the above reasons, a method for measuring the amount of gas ventilation sent (or excreted) into the patient's lung at one time by high-frequency ventilation has not been established.
[0011]
For this reason, when performing high-frequency ventilation, doctors do not have an index for managing the state of ventilation of the patient, and whether normal ventilation is being performed depends on the magnitude of the tidal volume and the chest of the patient. It was performed by visual observation, etc., with the degree of swinging being different as a guideline, and the accuracy was quite low, and the handling of the high frequency ventilator itself required skill and experience from the doctor .
[0012]
OBJECT OF THE INVENTION
The present invention provides a high-frequency ventilator capable of detecting the intrapulmonary pressure in order to detect the ventilation volume in the patient's lung during high-frequency ventilation, which improves the disadvantages of the conventional example. Objective.
[0013]
[Means for Solving the Problems]
According to the first aspect of the present invention, an inhalation introduction unit that supplies inhalation containing oxygen to the patient, an oscillating air pressure urging unit that urges the inspiration to oscillate air pressure having a period higher than the patient's respiratory cycle, and oscillating air pressure. A patient-side path for guiding the inhaled air to the patient, and an exhaust path for discharging exhaled air containing carbon dioxide from the patient to the atmosphere.
[0014]
And it equips with the patient side edge part of a patient side path | route the endotracheal insertion tube inserted from the patient's mouth to near the 1st branch of a trachea.
[0015]
Further, a detection branch pipe branched from any part of the discharge path from the patient side path, a pressure detection container communicating with any of the paths via the detection branch pipe, and an internal pressure of the pressure detection container And a corresponding pressure detection sensor. The inner diameter and length of the detection branch tube and the volume of the pressure detection container are related to the linear function of the pressure on the insertion end of the insertion tube and the internal pressure of the pressure detection container during high-frequency artificial respiration. It is set to a value that generates.
[0016]
In addition, a controller having a function of storing the relationship of the linear function and calculating the pressure at the distal end portion on the insertion side of the endotracheal insertion tube based on the detected pressure from the corresponding pressure detection sensor is provided.
[0017]
In the above configuration, first, the endotracheal tube is inserted from the patient's mouth to the first branch of the trachea. In such a state, intake air is supplied at a constant flow rate from the intake air introduction portion, and at the same time, the vibration air pressure urging portion is driven with a constant vibration cycle. Thereby, inhalation is supplied to a patient via a patient side course and an endotracheal insertion tube. Since the oscillating air pressure is applied to the inhalation, the inhalation air enters and diffuses into the lung by the positive pressure of the oscillating air pressure, and the exhalation is sucked out from the lung by the negative pressure. Such exhaled air reaches the discharge path via the endotracheal tube and is discharged into the atmosphere together with the subsequent inspiration.
[0018]
At this time, the corresponding pressure detection sensor detects the pressure in the pressure detection container and outputs it to the controller. In the high-frequency ventilator, the detection branch pipe and the pressure detection container are selected in which the insertion-side tip pressure of the endotracheal insertion pipe during high-frequency ventilation and the internal pressure of the pressure detection container have a linear function relationship. Since it is equipped, the controller uses the linear function to calculate the pressure at the distal end of the endotracheal tube, that is, the pressure in the patient's lung, from the output of the corresponding pressure detection sensor.
[0019]
In the second aspect of the invention, the controller has the same configuration as that of the first aspect of the invention, and the controller stores a linear function, and the pressure on the insertion-side tip of the endotracheal insertion tube with reference to the linear function. And an intrapulmonary pressure calculator that calculates the pressure of the lung. In the second aspect of the invention, the same operation as that of the first aspect of the invention is performed, and the calculation of the pressure (pulmonary pressure) at the distal end of the endotracheal tube is performed by the intrapulmonary pressure calculation unit. It is done with reference.
[0020]
The invention described in claim 3 has the same configuration as that of the invention described in claim 1 or 2, and has a configuration in which an operation panel for inputting a linear function from the outside to the storage unit is provided in the controller. In the invention described in claim 3, the same operation as that of the invention described in claim 1 or 2 is performed, and the pressure at the distal end portion of the endotracheal tube (intrapulmonary pressure) is calculated according to a linear function previously input from the operation panel. Is done.
[0021]
The invention described in claim 4 has the same configuration as that of the invention described in claim 1, 2 or 3, and a display unit for displaying the pressure of the insertion side distal end pressure of the endotracheal tube calculated by the controller to the outside. , Adopting the configuration of being attached to the controller. In such a configuration, operations similar to those in the above-described configurations are performed, and the calculated pressure at the distal end portion of the endotracheal tube (intrapulmonary pressure) is displayed on the display unit.
[0022]
The invention according to claim 5 has the same configuration as that of the invention according to claim 1, 2, 3 or 4, and the controller calculates the patient's pressure based on the calculated pressure at the distal end of the insertion side of the endotracheal insertion tube. A configuration is adopted in which a ventilation amount calculation unit for calculating a ventilation amount for the lung is provided. This ventilation volume calculation unit obtains a ventilation volume (how much inspiration is flowing into the lung or exhalation is exhausted) from the calculated intrapulmonary pressure.
[0023]
The present invention intends to achieve the above-described object by the above-described configurations.
[0024]
DETAILED DESCRIPTION OF THE INVENTION
(Overall configuration of the embodiment)
An embodiment of the present invention will be described with reference to FIGS. 1 and 2 are block diagrams showing the configuration of the high-frequency ventilator 12 according to the present embodiment, respectively. FIG. 1 shows a state during artificial respiration, and FIG. 2 is a preparatory work stage before performing artificial respiration. Shows the state.
[0025]
The high-frequency ventilator 12 includes an inhalation introduction unit 62 that supplies inhalation containing oxygen to the patient X, and an oscillating air pressure energizing unit that energizes the oscillating air pressure in a period higher than the respiration cycle of the patient X to the inhalation. 50, a patient-side path 60 that guides inhalation in which vibration air pressure is energized to the patient, a discharge path 70 that discharges exhaled air containing carbon dioxide from the patient X to the atmosphere, and operation control of the above parts The controller 40 which performs is provided.
[0026]
Each part is described in detail below.
[0027]
(Intake inlet)
The intake air introduction unit 62 includes a blender 621 that sucks and mixes outside air and oxygen prepared in advance, and a humidifier 622 that humidifies air sent from the blender 621. The blender 621 is provided with a plurality of output valves (not shown) that allow intake air to flow to the humidifier 622 side. The various output valves have different flow rates, and intake air at a predetermined flow rate is supplied by selecting an output valve with an arbitrary flow rate. Each output valve is provided with an actuator that switches between opening and closing by an operation signal from the controller 40.
[0028]
The humidifier 622 is connected to an intake pipe 623 that supplies the inhaled air Ai passed through the humidifier 622 to the patient X. The intake pipe 623 branches in the middle, and one end side communicates with a pressurized chamber 563 of a diaphragm mechanism 56 described later, and the other end side is connected to a three-way branch pipe 170 described later.
[0029]
(Vibration air pressure energizing part)
The oscillating air pressure urging unit 50 alternately selects a blower 52 that generates both positive pressure Ap and negative pressure An simultaneously, and a positive pressure Ap or negative pressure An generated by the blower 52, and selects a predetermined oscillating air pressure Apn. The rotary valve mechanism 54 that converts to the pressure, and the oscillating air pressure Apn from the rotary valve mechanism 54 actuates and operates to oxygen supplied to the patient X from the inhalation introduction part 62 (strictly, oxygen mixed with air). A configuration including a diaphragm mechanism 56 for energizing the oscillating air pressure is adopted.
[0030]
The blower 52 described above generates positive pressure and negative pressure at the same time by taking air into and sending out the air. The air intake port is connected to a negative pressure port 542 of a rotary valve mechanism 54 described later, and the air delivery port is connected to a positive pressure port 541.
[0031]
The rotary valve mechanism 54 includes a positive pressure port 541 to which a positive pressure is input from the blower 52, a negative pressure port 542 to which a negative pressure is urged from the blower 52, an output port 543 that outputs vibration air pressure, and its rotation. Thus, the output port 543 is composed of a rotary valve 544 that alternately connects the positive pressure port 541 and the negative pressure port 542, and a drive unit 545 that rotates the rotary valve 544.
[0032]
The drive unit 545 includes an electric motor and a speed reducer (not shown), and rotates the rotary valve 544 at, for example, 900 [rpm]. Each time the rotary valve 544 rotates, only the port 541 and the port 543 are communicated once, and then only the port 542 and the port 543 are communicated once. Thereby, the oscillating air pressure Apn having a frequency of 15 [Hz] is urged to the supplied intake air. An oscillating air pressure tube 546 that transmits the oscillating air pressure Apn to the diaphragm mechanism 56 is connected to the port 543.
[0033]
The diaphragm mechanism 56 includes a pressurizing chamber 562 and a pressurized chamber 563, and a diaphragm 561 formed of a stretchable film member that partitions the pressurized chamber 562 and the pressurized chamber 563. . The pressurizing chamber 562 is connected to the oscillating pneumatic tube 546. The pressurizing chamber 562 is connected to the output port 543 of the rotary valve 54, and the pressurized chamber 563 is connected to the intake pipe 623. With this structure, the oscillating air pressure formed by the rotary valve 54 is urged by the intake air flowing through the intake pipe 623 via the diaphragm 561.
[0034]
(Patient route)
Further, the high-frequency ventilator 12 includes a three-way branch pipe 170 on the downstream side of the intake pipe 623, and the three-way branch pipe 170 further branches the downstream side into the patient X side and the discharge path side. The three-way branch pipe 170 includes three pipe lines, a patient-side pipe line 171 (patient-side end), an oxygen supply source-side pipe line 172, and an exhalation-exhaust side pipe line 173, all of which are internal. It is joined at. The oxygen supply source side pipe line 172 is connected to the inspiratory pipe 623, and the patient side pipe line 171 is connected to the endotracheal insertion pipe 81 reaching the lungs of the patient X. The three-way branch pipe 170 and the intake pipe 623 constitute a patient side path 60. In addition, a patient-side pressure sensor 93 that detects an average airway pressure is provided in the patient-side conduit 171, and the detected pressure is output to the controller 40.
[0035]
(Discharge route)
Further, the exhalation discharge side pipe line 173 of the three-way branch pipe 170 is connected to one end portion of the discharge pipe 604, and a flow rate adjusting valve 607 is connected to the other end portion of the discharge pipe 604. The exhaust pipe 604 and the flow rate control valve 607 serve as a passage for inspiration (exhalation) containing carbon dioxide emitted from the lungs of the patient X, and constitute an exhaust path 70 for exhausting the exhalation into the atmosphere.
[0036]
FIG. 3 is an enlarged view in which the periphery of the discharge path 70 is partially cut away. As shown in this figure, the flow rate adjusting valve 607 includes a housing 607a, an exhaust port 607b, a flow control valve (control silicon sheet) 607c, and a reciprocating movement that moves the transfer valve 607c back and forth along a certain direction. And a solenoid 607d as an urging mechanism.
[0037]
(Model lung and endotracheal tube)
By the way, as can be seen by comparing FIG. 1 and FIG. 2, in the preparatory work stage before artificial respiration (FIG. 2), the patient side conduit 171 of the three-way branch pipe 170 is connected to the trachea from the mouth of the patient X. A model lung 82 is provided which is set to a compliance value approximately equal to the lungs of patient X via an endotracheal tube 81 which can be inserted up to one branch.
[0038]
The endotracheal insertion tube 81 is used not only during the preparatory work stage but also during artificial respiration, and is inserted into the trachea from the mouth of the patient X during the artificial respiration. The human trachea diverges into two bronchi in the deep part, each going to the left and right lungs. This endotracheal insertion tube 81 is tubular and highly deformable, and is inserted from the mouth of the patient X to the above-mentioned bronchus (first branch) during artificial respiration as shown in FIG. The Accordingly, the endotracheal insertion tube 81 is set to a length that can be sufficiently reached from the mouth of the patient X to the first branch, and of course, is set to an outer diameter that can be inserted into the trachea.
[0039]
The endotracheal tube 81 is not configured so that each part is dimensioned according to a specific patient X alone. That is, there is no great difference in the inner diameter and length of the trachea for general adults, and the dimensions are set for such average adults. That is, in the case of an adult male, the length from the mouth to the first branch is about 22 to 26 [cm], and the length from the patient side conduit 171 of the three-way branch pipe 170 to the mouth is +3 to 5 [cm]. Therefore, the total length of the endotracheal tube 81 may be about 25 to 31 [cm], and is set to 30 [cm] in the present embodiment. When a normal adult is a subject, a tube having an inner diameter of the endotracheal tube 81 of about 8 [mm] is generally used.
[0040]
The dimensions of each part of the endotracheal tube 81 may be set to a size according to the age of the patient X. FIG. 4 shows the age and the inner diameter and insertion length of the endotracheal tube according to the age. Since the insertion length indicates the length of the portion inserted from the mouth, the actual length of each endotracheal insertion tube is obtained by adding 3 to 5 [cm] to this value.
[0041]
Further, the endotracheal tube 81 is replaceable, and is detachable from the patient-side conduit 171 of the three-way branch tube 170. Therefore, after being used for artificial respiration, it is removed and discarded or sanitized and reused.
[0042]
In the preparatory work stage prior to artificial respiration, a model lung 82 is provided at the end of the endotracheal tube 81. The model lung 82 has a hollow inside, and its outer shape is formed in a cylindrical shape having a cylindrical shape. In addition, a connection port with the endotracheal insertion tube 81 is provided in the upper part, and the inside of the model lung 82 communicates with the endotracheal insertion tube 81. The model lung 82 is set to be approximately equal to the lung compliance value of the patient X. Here, strictly speaking, compliance means an index indicating the ease of extension of an object. However, here, the compliance (C) indicates the inflow or outflow amount (ΔV) of gas into the container with respect to a constant pressure change (ΔP) (C = ΔV / ΔP; unit [ml / cmH 2 O ]).
[0043]
Human lung compliance depends on lung elasticity and volume. On the other hand, the model lung 82 is formed of a hard material with poor stretchability, and its compliance is adjusted to be equal to the lung compliance of the patient X by appropriately setting only the volume thereof. are doing. This is because it is difficult to adjust the compliance by forming the model lung 82 with a material having elasticity similar to that of the lung from the viewpoint of productivity, and it is more productive to adjust the compliance only from the volume setting. This is because it becomes easy.
[0044]
In the present embodiment, the volume of the model lung 82 is set to 40 [l] based on the general compliance value of a normal adult. Further, since the compliance of the lung can be measured by a dedicated measurement device, the compliance may be measured for each patient in advance, and the volume of the model lung 82 may be determined more strictly based on the measured value.
[0045]
Further, the model lung 82 is equipped with a model lung pressure sensor 83 for detecting the pressure inside the model lung 82. The model lung pressure sensor 83 is connected to an observation apparatus B that stores and displays the sensor output. This observation apparatus B is composed of, for example, a personal computer and a display that are separate from the high-frequency ventilator 12.
[0046]
The model lung 82 is detachable from the endotracheal insertion tube 81, and is removed from the endotracheal insertion tube 81 when the preparatory work before artificial respiration described later is completed.
[0047]
(Pressure detection container and branch pipe for detection)
As shown in FIGS. 1 and 2, the high-frequency ventilator 12 is further connected to a detection branch pipe 84 branched from the vicinity of the end of the intake pipe 623 on the three-way branch pipe 170 side. A pressure detection container 85 communicating with the intake pipe 623 is provided. The detection branch tube 84 is a tube-shaped tube having an inner diameter smaller than that of the aforementioned endotracheal tube 81 (for example, 0.5 [mm]) and a length of 30 [cm]. The detection branch pipe 84 is detachably attached to the intake pipe 623.
[0048]
Further, the pressure detection container 85 is hollow inside, and its outer shape is formed in a cylindrical shape with a cylindrical shape. In addition, a connection port with the detection branch pipe 84 is provided in the upper part thereof, and the inside of the pressure detection container 85 communicates with the intake pipe 623 through the detection branch pipe 84. The pressure detection container 85 is detachably connected from the detection branch pipe 84.
[0049]
The pressure detection container 85 is not set to be equal to the compliance value of the lung of the patient X like the model lung 82 but is made of a hard material like the model lung 82. Further, the volume of the pressure detection container 85 is set to 10 [ml] which is smaller than the model lung 82 described above.
[0050]
The values of the inner diameter and length of the detection branch pipe 84 and the volume of the pressure detection container 85 are the values of high-frequency artificial respiration (intake supply amount, vibration cycle of vibration air pressure and vibration air pressure with the model lung 82 connected). Located at the distal end on the insertion side of the endotracheal tube 84 when the urging unit 50 performs a single ventilation amount (under a condition where each value of the oscillating air pressure in one cycle) is a constant value. The internal pressure P1 of the model lung 82 and the internal pressure P2 of the pressure detection container 85 are always set to values that cause a linear function relationship.
[0051]
This linear function is expressed by p1 = k · p2 + m (k and m are constants), and the values of k and m are the dimensions of each part of the endotracheal tube 81, model lung 82, detection branch tube 84, and pressure detection container 85. Or it is determined according to the volume. Further, the values of the inner diameter and length of the detection branch tube 84 and the volume of the pressure detection container 85 are repeatedly changed in a state in which the model lung 82 is connected to the endotracheal tube 81 with the values changed. An artificial respiration test is performed and the one showing the most linear change is adopted. Incidentally, each numerical value described above is a value obtained by a test.
[0052]
When the inner diameter of the detection branch pipe 84 is small or long, the volume of the pressure detection container 85 tends to be small. Accordingly, the inner diameter and length of the detection branch pipe 84 and the volume of the pressure detection container 85 are not particularly limited to this volume. In other words, even if the inner diameter and length of the endotracheal tube 81 and the volume of the model lung 82 are constant, there are innumerable combinations of the inner diameter and length of the detection branch tube 84 and the volume value of the pressure detection container 85. Set to one of these.
[0053]
Further, the pressure detection container 85 is equipped with a corresponding pressure detection sensor 86 for detecting the pressure inside the container. The corresponding pressure detection sensor 86 is connected to a storage unit 41 of the controller 40 described later during high-frequency artificial respiration, and is connected to the observation device B described above in a preparatory work stage before performing artificial respiration.
[0054]
(controller)
Next, the controller 40 will be described with reference to FIGS. FIG. 5 is a block diagram showing a control system of the high frequency ventilator 12. The controller 40 is composed of an arithmetic unit including a CPU, a ROM, and an A / D converter, and a program for executing operation control of the high-frequency ventilator 12 described later is input. The controller 40 controls the detected pressure of the corresponding pressure detection sensor 86 and the operation control unit 49 that controls the operation of the intake air introduction unit 62, the oscillating air pressure urging unit 50 and the flow rate adjusting valve 607 in accordance with the input conditions of the operation panel 43 described later. The sensor output memory 41 to be stored, the storage unit 42 to store the above-described linear function, the storage unit 42 and the sensor output memory 41 are referred to, and the model lung pressure sensor 83 uses the detected pressure of the corresponding pressure detection sensor 86. The intrapulmonary pressure calculation unit 47 that calculates the pressure that will be detected (the intrapulmonary pressure of the patient X determined by the calculation), and the ventilation amount for the lungs of the patient X based on the output of the intrapulmonary pressure calculation unit 47 A ventilation amount calculating unit 48 for calculating is provided.
[0055]
Further, the controller 40 is provided with an operation panel 43 for inputting each operation described later and a display unit 44 for displaying the pressure specified by the patient side pressure specifying unit 87.
[0056]
The operation control of the controller 40 will be described in detail. In this high-frequency ventilator 12, in order to enable observation of the intake and exhaust ventilation based on the oscillating air pressure in the lungs of the patient X during the high-frequency ventilation, the high-frequency ventilator 12 The purpose is to enable observation of pressure.
[0057]
In order to achieve such an object, the above-described high-frequency ventilator 12 requires preparatory work before performing artificial respiration. This preparatory work is detected by the model lung 82 and the pressure detection container 85 under the condition that the high-frequency artificial respiration is supplied and supplied, the vibration cycle of the oscillating air pressure, and the tidal volume in the oscillating air pressure urging unit 50 are constant. This is performed in order to select the inner diameter and length of the detection branch pipe 84 and the volume of the pressure detection container 85 so that each pressure has a corresponding relationship represented by a linear function.
[0058]
This will be explained in order. First, the endotracheal tube 81 to be inserted into the trachea of the patient X is connected to the model lung 82, and the model lung 82 and the patient side conduit 171 of the three-way branch tube 170 are connected. Communicate. The model lung pressure sensor 83 and the corresponding pressure detection sensor 86 are connected to the observation apparatus B.
[0059]
Then, the operation panel 43 is used to input and set the intake air flow rate, the vibration period of the oscillating air pressure and the tidal volume of the oscillating air pressure urging unit. When such an input is made, the operation control unit 49 selects an appropriate one from the output valves provided in the blender 621 of the intake air introduction unit 62, and at the same time, the rotary valve mechanism 54 of the oscillating air pressure urging unit 50 is input. The blower 52 is driven at an output corresponding to the tidal volume input. As a result, inhalation is supplied to the model lung 82 and the pressure detection container 85. At this time, the opening degree of the flow rate adjustment valve 607 is adjusted from the operation panel 43 via the operation control unit 49 so that the inspiratory pressure becomes a numerical value optimum for the patient.
[0060]
Each pressure based on the outputs of the model lung pressure sensor 83 and the corresponding pressure detection sensor 86 observed at this time is compared, and it is verified whether or not the correspondence relationship of the linear function described above has occurred. Since each detected pressure repeats increasing / decreasing at the same period as the oscillating air pressure, the detected pressure of any one of the sensors 83 and 86 is selected, and the detected pressures at the same phase are compared.
[0061]
When the correspondence relationship of the linear function described above is not generated, the detection branch pipe 84 is replaced with one having a different inner diameter size. Alternatively, the pressure detection container 85 is replaced with a different volume. When appropriate detection branch pipe 84 and pressure detection container 85 are specified by such detection, linear functions constants k and m are calculated from sensors 83 and 86 described above. Then, the calculated linear function is input from the operation panel 43.
[0062]
Next, a case where high-frequency artificial respiration is performed on the patient X will be described. First, the model lung 82 is removed from the endotracheal tube 81, and the endotracheal tube 81 is inserted from the mouth of the patient X to the first branch of the trachea. In addition, the corresponding pressure detection sensor 86 is connected to the controller 40. Then, the operation panel 43 inputs the lung compliance value of the patient X and the same inspiratory flow rate, vibration frequency of the oscillating air pressure and tidal volume as in the preparation stage.
[0063]
As a result, the intake air is supplied at a constant flow rate selected from the intake air introduction unit 62, and at the same time, the rotary valve mechanism 54 of the oscillating air pressure urging unit 50 is driven at the rotational speed corresponding to the input vibration cycle, and the blower 52 is input. Drive with the output corresponding to the tidal volume. As a result, the oscillating air pressure is urged at the same period as that during the preparatory work, and the inspiratory air urged by the oscillating air pressure is supplied to the lungs of the patient X and the pressure detection container 85.
[0064]
At this time, the intrapulmonary pressure calculation unit 47 of the controller 40 substitutes the detected pressure p2 output from the corresponding pressure detection sensor 86 into the above-described relative relational expression p1 = k · p2 + m, and the model lung 82 is connected. For example, the pressure that would have been detected by the model lung pressure sensor 83 is calculated. Since the compliance of the model lung 82 is set equal to the compliance of the lung of the patient X, the pressure that would be detected by the model lung pressure sensor 83 calculated by the intrapulmonary pressure calculation unit 47 is the same as the intrapulmonary pressure of the patient X. Almost equal. Therefore, this calculated pressure is output on the display unit 44 as the current intrapulmonary pressure of the patient.
[0065]
Further, in the ventilation volume calculation unit 48, ΔV = k · Δp2 · C formula (ΔV: ventilation volume (inflow or outflow rate of gas to the lung in the atmospheric pressure state), k: coefficient obtained by the relative relationship calculation unit 46, Δp2: Pressure change (atmospheric pressure—detected pressure p2 of the corresponding pressure detection sensor), C: lung compliance value) is calculated and output on the display unit 44.
[0066]
With the above configuration, the high-frequency ventilator 12 can simultaneously detect the intrapulmonary pressure and the ventilation volume of the patient X during the high-frequency ventilation while performing high-frequency ventilation.
[0067]
(Principle of this embodiment)
〔compliance〕
First, the hydrodynamic correlation between the “real lung” and the “model lung 82” when performing high-frequency ventilation (HFO) will be described. “Actual lung” is a non-linear elastic body. When inhalation supply and exhalation are repeated at a period close to that of spontaneous respiration as in normal artificial respiration, “internal pressure-ventilation” depends on inspiration supply and exhalation. The curve of “amount” is different, and there is hysteresis shown in FIG.
[0068]
In addition, as shown in the lung model shown in FIG. 7, the “actual lung” exists at the end of each of the bronchi and the innumerable bronchi that have been branched multiple times from a single trachea. Consisting of countless small alveoli.
[0069]
The average pressure value of the pressure applied to all alveoli is the representative value P of the patient's intrapulmonary pressure, and the difficulty of inhalation inflow into the lungs is the airway resistance R (typical air resistance is the trachea), When compliance is considered compliance C, the change in intrapulmonary pressure associated with actual gas ventilation V is generally expressed by the following equation from various papers that have studied medical mechanics: It has become.
[0070]
ΔP = R · (dΔV / dt) + ΔV / C (1)
(ΔP: patient's intrapulmonary pressure change [cmH2O], R: airway resistance until flowing into the alveoli [cmH2O / l / sec], C: compliance [ml / cmH2O], ΔV: change in ventilation [ml])
[0071]
Here, the definition of the compliance C indicates the inflow or outflow amount (ΔV) of gas into the container with respect to a constant pressure change (ΔP) as described above, and C = ΔV / ΔP (unit: [ml / cmH2O]). The airway resistance R is a frictional resistance generated between gas molecules or between the gas molecules and the airway wall surface when the gas flows in the airway, and is a pressure required to flow the gas at a constant flow rate (the unit is [ cmH2O · s / l]).
[0072]
The above-described specific method for changing the internal pressure with respect to the lungs during artificial respiration in the high-frequency ventilator 12 models the trachea and lungs of the patient X so that the flow resistance and the compliance are equal, and the model (intratracheal insertion) is preliminarily modeled. By performing high-frequency artificial respiration on the tube 81 and the model lung 82), the length and inner diameter of the detection branch tube 84 and the volume of the pressure detection vessel 85 are obtained as appropriate, and the relative primary relationship is obtained. The function constants k and m are calculated and input to the storage unit 42 of the controller 40. Then, in the lung pressure calculation unit 47 of the controller 40, the actual change in the internal pressure of the lung is specified only from the change in the internal pressure observed in the pressure detection container 85 from the above-described relative relationship. In addition, by specifying the actual change in the internal pressure of the lung in this way, it is possible to specify the lung ventilation during artificial respiration according to equation (16) described later.
[0073]
However, in the actual lung, as shown in FIG. 6, it is expected that it is difficult to specify the compliance C because hysteresis occurs during intake and exhaust. In addition, it is expected that it is difficult to obtain the flow resistance because of the structure that repeats countless branches from the trachea to the alveoli. Therefore, actual lung and trachea modeling may be difficult.
[0074]
However, in the case of high-frequency ventilation, the “model lung 82” can be used as a lung substitute model because of the following unique ventilation method that is different from the normal ventilation method.
[0075]
(1) Since an average pressure is always applied to the lungs (5 to 15 [cmH20]), the lungs are always inflated to some extent. As a result, the lungs move less.
[0076]
(2) Compared to the normal and large and slow ventilation (tidal volume = 500 to 700 [ml], ventilation rate = 8 to 20 [t / min]), the tidal volume is very small ( 100 [ml] or less), ventilation rate is very fast (180 [times / min] or more). Therefore, the movement of the lungs due to ventilation is small.
[0077]
(3) In normal ventilation, the instrument actively pushes in at the time of inhalation, and at the time of exhalation, the instrument does not participate in the exhalation of gas and is a flow that naturally exhales using the elasticity of the lungs. Largely affected by the elasticity of the lungs. Therefore, hysteresis appears as shown in FIG. 6 “pulmonary pressure-ventilation curve”. However, in the case of gas ventilation for high-frequency ventilation, the instrument performs active and forced quantitative single ventilation within a fixed time regardless of the elasticity of the lungs, both at the time of inhalation supply and at the time of exhalation. . As a result, there is no hysteresis in lung motion.
[0078]
For the reasons of (1), (2), and (3), during high-frequency artificial respiration, the movement of the lungs during normal ventilation, “lung expansion (inhalation)-lung contraction (in expiration)” is small. Become. That is, there is no noticeable movement of the lung made of a non-linear elastic body due to breathing, and the difference between the compliance C and the exhaust time becomes small. For this reason, the modeling of the lung equal to the compliance C during high-frequency artificial respiration can be approximately realized by the “model lung 82” which is a sealed single volume tank with low elasticity.
[0079]
The overall compliance of the human lung includes static compliance, dynamic compliance, and thoracic septal compliance. Since normal ventilation has a large respiratory movement (inspiratory-exhalation), it is necessary to consider the effects of dynamic lung compliance and thoracic compliance in consideration of lung hysteresis. Because the change in volume (ventilation volume) occurring in the lung is small in terms of time and volume due to the law, the dynamic lung compliance and the thoracic compliance, which are affected dynamically by the lung, are small, You can ignore it.
[0080]
In the case of high-frequency ventilation, there is no problem with the static compliance of the lung as the overall compliance of the lung, so that value is used as a representative value. The static compliance C of medically normal lung is about 200 [ml / cmH20]. Therefore, in the case of an adult, a circuit using “model lung 82” having compliance C based on the following equation (2) may be used as a reference.
[0081]
ΔP = R · (dΔV / dt) + ΔV / 200 (2)
[0082]
[Airway resistance]
Also consider airway resistance R. FIG. 8 shows an “airway bifurcation configuration”. The lung is composed of many branches from the trachea to the alveoli. Assuming that the trachea is 0-dimensional and the order is added for each branch, the final alveolar dimension is 20 or more.
[0083]
The airway diameter becomes narrower as the number of airway branches increases. However, the number of airways increases conversely. FIG. 9 shows the relationship between the bifurcation dimension and the sum of the airway cross sections for each dimension. According to this, the sum of the airway cross-sectional areas (the sum of the cross-sectional areas of all airways of the same dimension) remains almost unchanged until the tenth branch. However, after the tenth branch, the airway is divided innumerably and the airway cross-sectional area increases rapidly.
[0084]
Airway resistance R increases in inverse proportion to the fourth power of the airway radius from the Poiseui 11e equation, but conversely, the sum of the airway cross-sectional areas increases rapidly. The resistance R becomes small. The Poiseuille equation is shown below.
[0085]
R = 8 nl / πr ⁴ ... (3)
(N: gas viscosity coefficient (poises), l: length [cm], r: radius [cm])
[0086]
FIG. 10 shows the relationship between the branching dimension and the airway resistance due to the trachea in each dimension. According to this, the majority of the airway resistance R is the airway up to the middle-sized bronchus (seventh branch) of the lung, which produces a resistance of 80% or more. Considering up to the 10th branch, it is over 95%. Therefore, the airway resistance R may be considered as a resistance from approximately the 0th branch to the 10th branch.
[0087]
Here, in normal breathing, branch 0 is the trachea. However, in the case of a patient during high-frequency artificial respiration, since the endotracheal tube 81 is inserted in the 0 branch of the patient, the airway resistance of the 0 branch may be considered as the resistance of the endotracheal tube 81. That is, the air resistance of the 0 branch may be considered to be almost the same as the resistance generated by substituting the inner diameter of the endotracheal insertion tube 81 into the Poiseuille equation.
[0088]
Next, the resistance Rc of the endotracheal tube 81 is compared with the magnitude of the first branch resistor R1 + second branch resistor R2 + third branch resistor R3 +... + Tenth branch resistor R10. Assuming that the diameter of the endotracheal tube 81 is 8 [mm] and the tube length is 30 [cm], the resistance Rc of the endotracheal tube 81 is expressed by the following equation (4).
[0089]

(N: viscosity coefficient of gas, r: radius, I: length)
[0090]
The first branch resistor R1 +... + The tenth branch resistor R10 is considered as follows. According to the diagram shown in FIG. 9, the airway cumulative area (the sum of the cross-sectional areas of all airways of the same dimension) of each branch from the first to the tenth branches is substantially equal. Although it varies depending on the patient, a representative example of the accumulated airway area from the first to the fourth branch of the human is referred to the values in FIG. The average cross-sectional area of the airway from the first to the fourth branch is 46 [cm ² It can be about. If this is almost the same cumulative area from the first to the tenth branch, it is assumed that the first to the tenth branch has an air resistance equivalent to that of a single tube of 3.8 [cm] in diameter. Good. Assuming that the length of the first to tenth branches is about 20 [cm], the total airway resistance from the first branch to the tenth branch in the lung can be expressed by the following equation (5).
[0091]
1st branch resistor R1 + ... + 10th branch resistor R10
= (8 × n × l) / πr ^Four = (8 × n × 20) / (π · 1.4 ^Four ) = 41.6n / π = 13.3n (5)
[0092]
Since the viscosity coefficient n of the gas is constant, the “intratracheal tube 81 resistance Rc” is 2984n / 13.3n = about 224 times the “first branch resistance R1 +... + 10th branch resistance R10”. .
[0093]
Therefore, since the intratracheal insertion tube 81 resistance Rc >> the first branch resistance R1 +... + The tenth branch resistance R10, the airway resistance R is almost generated by the endotracheal insertion tube 81 resistance Rc, and may be considered to be substantially equivalent. In particular, breathing methods that produce fast and small flow, such as high-frequency ventilation, do not directly reach the deep branch, so the actual air flow is mainly due to the low order. Up to the branches. Therefore, the term of R that should be considered when there is an air flow has a stronger tendency of “total resistance in the tracheostomy tube 81 Rc >> first branch and subsequent branches”. In the high-frequency artificial respiration method, ventilation is performed by diffusion of oxygen rather than spreading oxygen by direct air flow, so ventilation is sufficient even when direct flow does not reach deep branches. To be done.
[0094]
As described above, when the high-frequency artificial ventilation method is employed, the airway resistance R is represented by the air resistance of the “intratracheal insertion tube 81” to be used.
[0095]
For these reasons, the R term of ΔP = R · ΔV / dt + ΔV / C in equation (1) is the airway resistance Rc, which is the air resistance of the “intratracheal tube 81”, and the C term is about 200 [ml / cmH20]. By setting the compliance C of the “model lung 82” of the hard tank, a simple model can be obtained.
[0096]
Using these modeled HFO circuits, “pressure detection container 85” and “changes based on the number of hours described above” that are the same as or similar to changes in pressure P1 occurring in “model lung 82” and “ If the HFO respiratory device having the same circuit configuration is used for the patient, the change in the pressure P3 of the “actual lung” indirectly results in the pressure P2 of the “pressure detection container 85”. You can know by measuring the change of.
[0097]
The above can be summarized as follows. When the change in the lung pressure of the patient at the time of HFO is ΔP, the change amount (flow rate) of the gas entering and leaving the lung is ΔV, the airway resistance of the lung at the time of HFO is R, and the compliance of the lung is C (1) ) Holds.
[0098]
At this time, as described above, during the high-frequency artificial respiration, there is almost no problem even if the term R is approximately replaced with the air resistance Rc of the “intratracheal insertion tube 81” used for the patient. Furthermore, since the compliance C can be the compliance of the model lung 82 (= 200 [ml / cmH 2 O]), the following equation (6) is established.
[0099]
ΔP = Rc · (dΔV / dt) + ΔV / 200 (6)
[0100]
[Pressure detection container and branch pipe for detection]
Here, the “branch tube 84 for detection” having the air resistance of Rd and the “pressure detection container 85” having the compliance of C ′ are expressed as “relationship between the model lung 82 and the pressure detection container 85” shown in FIG. Use in the same HFO circuit. However, it is assumed that Rd is in a relative relationship with Rc and C ′ is in a relative relationship with C. When the volume change occurring in the “pressure detection container 85” is ΔV ′, the following equation (7) is established.
[0101]
ΔP = Rd · (dΔV ′ / dt) + ΔV ′ / C ′ (7)
[0102]
The pressure change P1 occurring in the “model lung 82” changes the pressure according to the above-described equation (6), and the pressure change P2 occurring in the “pressure detection container 85” changes according to the above equation (7). Since all the terms in the above equation (7) are in a proportional relationship with the previous equation (6), the relationship between P1 and P2 is as follows.
[0103]
P1 = k × P2 + m (8)
(K and m are constants)
[0104]
Here, if the change in internal pressure of the patient's lung during HFO is Palv, the actual lung can be simply modeled as “model lung 82” from the ventilation mechanism peculiar to HFO. It can represent with following Formula (9).
[0105]
Palv≈P1 (9)
[0106]
Therefore, if the high-frequency ventilator 12 to which the actual patient is connected obtains the pressure change P2 of the “pressure detection container 85” in the state shown in FIG. 2 (artificial respiration mode), the patient is expressed by the equations (8) and (9). The internal pressure change Palv of the lung is determined.
[0107]
Palv≈P1 = k × P2 + m (10)
[0108]
The patient's intrapulmonary pressure change can be indirectly measured by the above equation (10). Here, the installation positions of the pressure detection container 85 and the detection branch pipe 84 are preferably as close to the model lung 82 and the endotracheal insertion pipe 81 as possible, but may be other positions.
[0109]
The above has considered that the static compliance C of a normal lung model is about 200 [ml / cmH20], and “model lung 82” is considered as a representative value. It can be expressed by applying a coefficient n. If the compliance of the pathological lung is C ″, it can be expressed by the following equation (11).
[0110]
C ″ = nC (11)
(N is a constant)
[0111]
Therefore, the internal pressure change of the pathological lung is transformed into the following equation (12) by transforming the equation (6).
[0112]
ΔP = Rc · (dΔV / dt) + ΔV / 200n (12)
[0113]
Since the second term of equation (12) is a term that is proportional to C, equation (10) is as follows.
[0114]
Palv≈P1 = k ′ × P2 + m (13)
(K 'is a proportional multiple constant of k considering the coefficient n)
[0115]
The relationship between k and k ′ is proportional to the relationship between the “model lung 82 ′” having the static compliance C ″ of the pathological lung and the “model lung 82” having the normal lung compliance. Is guided. Therefore, in actual clinical practice, when the static compliance C ″ of the pathological lung is measured in advance at the time of medical treatment of the patient and the relationship between the value and the normal lung is obtained by the equation (11), which k ′ is used. I know what to do. If this k ′ is used, the change P1 in the lung pressure of the patient can be obtained from the equation (13) from P2 even in the diseased lung.
[0116]
[Calculation of tidal volume]
Furthermore, it is possible to estimate how much tidal ventilation enters and leaves the actual patient's lung from the following calculation formula.
[0117]
As a reference, a change in pressure when a fixed amount of gas 100 [ml] is momentarily pushed into the “model lung 82” is ΔPa (constant) (ΔPa = 100 / C, for example, C = 200 [ml / cmH 2 O] If there is ΔPa = 0.5 [cmH2O]). If a pressure amplitude of ΔP1 is generated in the “model lung 82” by the ventilation of the high-frequency artificial respiration method, the relationship of the following equation (14) is established using Boyle's law.
[0118]
ΔPa / 100 = ΔP1 / ΔV = 1 / C (14)
[0119]
Further, since the pressure amplitude change ΔP2 of P2 of the pressure detection container at the time of high-frequency artificial respiration has a relationship of ΔP1 and the equation (10), the following equation (15) is established.
[0120]
ΔPa / 100 = ΔP1 / ΔV = (k · ΔP2 + m) / ΔV = 1 / C (15)
[0121]
Therefore, ΔV (ventilation volume) that is a volume change at the time of HFO of “model lung 82” is obtained.
[0122]
ΔV = 100 × (k · ΔP2 + m) / ΔPa = C · (k · ΔP2 + m) (16)
[0123]
Since the “model lung 82” and the “actual lung” are substantially the same in terms of fluid dynamics, the actual patient's lung ventilation can be approximately ΔV. That is, if the coefficients k and m are obtained by the measurement in the preparatory work stage and the pressure change ΔP2 is measured during the high-frequency artificial respiration, the ventilation volume of the patient's lung can be almost estimated using the equation (16).
[0124]
It should be noted that the compliance value C has an influence on the pressure detection of the model lung 82 and the pressure detection container 85 in comparison with the airway resistance R when detecting pressure during high-frequency artificial respiration, for the reasons (1) to (3) described above. It can be said that the main specific factors of the change in the internal pressure depend on the airway resistance of the endotracheal insertion tube 81 and the detection branch tube 84. Therefore, when setting the volume of the model lung 82, even if there is a slight error (± 150 [%]) from the most preferable volume calculated based on the compliance value C, a large error does not occur in the detected pressure (more accurate). It is of course desirable to set the most suitable value in order to increase the value). However, the compliance C must be accurate when calculating the ventilation amount ΔV from the equation (16).
[0125]
[First embodiment]
The result of measuring the relationship of equation (8) using the above-described system of the high-frequency ventilator 12 will be described (see FIG. 2).
[0126]
First, high-purity oxygen (about 100%) and air are mixed in the intake air introduction unit 62 to create a mixed gas of high-concentration oxygen having a value set by the doctor. At the same time, inspiratory gas (fresh gas) having a flow rate set by the doctor (10 to 60 [l / min]) is allowed to flow in a steady flow in the patient circuit. Vibration is generated from the oscillating air pressure urging unit 50 to the steady flow to generate vibration waves of air. The air vibration wave with the flow exchanges the lungs of the patient X.
[0127]
In addition, by appropriately controlling the outlet open area of the discharge path 70 with the flow rate adjustment valve 607, an appropriate average pressure (usually referred to as average airway pressure, which is set by a doctor so as not to place a burden on the patient's lungs. Average airway pressure) Pmean = 5-15 [cmH20]) is maintained in the circuit, and ventilation is performed while applying a certain positive pressure to the patient's lungs. FIG. 11 shows a typical pressure waveform example of the “model lung 82” in this preparatory work stage.
[0128]
Further, actual experimental results comparing the pressure changes of P1 of “model lung 82” and P2 of “pressure detection container 85” will be described. As experimental conditions, “model lung 82” and “intratracheal tube 81” when an adult patient is assumed are used. A “pressure detection vessel 85” and a “detection branch tube 84” having airway resistance and compliance relative to the “model lung 82” and the “intratracheal insertion tube 81” are made as an experiment.
[0129]
Experimental conditions were set as follows.
“Model lung 82” ……………… 40 [l] tank
“Intratracheal tube 81” ………… Tube 81 with an inner diameter of 8 [mm] (length 30 [cm])
"Pressure detection container 85" ………… 10 [ml] container
“Branch tube 84 for detection” ………… Tube 81 with an inner diameter of 0.5 [mm] (length 30 [cm])
Under the above experimental conditions, the experiment was performed by setting the frequency of the oscillating air pressure to 5 [Hz] and 15 [Hz]. The pressure waveform data of 5 [Hz] is shown in FIG. 12, and the experimental results are shown in FIG. Further, the pressure waveform data of 15 [Hz] is shown in FIG. 14, and the experimental results are shown in FIG.
[0130]
The relationship between P1 and P2 is stable in a relative relationship in any period, and can be expressed by the previous equation (8).
[0131]
[Second embodiment]
In addition, it becomes a problem how the doctor selects the corresponding “model lung 82 + intratracheal insertion tube 81” in order to measure the lung of the patient performing HFO ventilation.
[0132]
As described above, the size (inner diameter) and intubation depth (tube length) of the endotracheal insertion tube 81 are determined by the age of the patient in the “intratracheal insertion tube 81” used for the patient (see FIG. 4). This is generally treated as an index value based on medical statistical data used when performing artificial respiration management.
[0133]
Moreover, if the patient's age is determined, lung compliance is almost calculated from medical statistical data. Depending on the pathological condition of the lung, the compliance of the diseased lung is almost calculated from medical statistical data. In addition, there are devices that can be measured clinically, so they are easily required.
[0134]
Here, the experimental conditions are “model lung 82 ′” and “intratracheal tube 81 ′” when a child patient (about 3 years old) is assumed. A “pressure detection vessel 85 ′” and a “detection branch tube 84 ′” having airway resistance and compliance relative to the “model lung 82 ′” and “intratracheal insertion tube 81 ′” are manufactured as a prototype.
[0135]
Experimental conditions were set as follows.
"Model lung 82 '" ……………… 20 [l] tank
“Intratracheal tube 81 ′” ………… Inner diameter 4 [mm] Intratracheal tube 81 (Length 19 [cm])
"Pressure detection container 85 '" ... 12 [ml] container
“Branch tube for detection 84 ′” ………… Inner diameter 0.5 [mm] Endotracheal tube 81 (Length 120 [cm])
Under the above experimental conditions, the experiment was performed by setting the frequency of the oscillating air pressure to 5 [Hz] and 15 [Hz]. The experimental result of 5 [Hz] is shown in FIG. Moreover, the experimental result of 15 [Hz] is shown in FIG.
[0136]
The relationship between P1 and P2 is stable in a relative relationship in any period, and can be expressed by Expression (8) ′ similar to the previous Expression (8).
[0137]
P1 ′ = k ′ × P2 ′ + m ′ (8) ′
(K 'and m' are constants)
[0138]
As described above, in the actual experimental results, changes in the patient's intrapulmonary pressure from any model patient's lung with the model lung having the same compliance value as the patient's lung and the pressure detection container similar to this model lung. Can be known relatively.
[0139]
The experimental results also show that the intrapulmonary pressure can be measured accurately by taking into account the coefficient multiplied by P2 according to the patient's medical condition.
[0140]
Therefore, since the dimensions of each part of the “intratracheal tube 81” and “patient lung compliance” are both arbitrarily determined depending on the age of the patient, these values may be handled as a set of data combined by age. . In addition, the length and inner diameter dimensions of the “branch tube 84 for detection” and the value of “compliance of the pressure detection container 85 (or the volume of the pressure detection container 85)”, which are relatively matched with each other, are also set according to age. May be handled as
[0141]
As a result, as long as the doctor decides the endotracheal tube 81 to be used for the patient receiving HFO, the doctor can also set the “pressure detection vessel 85” and the “branch tube 84 for detection” for measuring the intrapulmonary pressure. Since it is decided at the same time without hesitation, it is easy to select and easy to handle. In addition, the presence of the “pressure detection container 85” and the “branch tube 84 for detection” in accordance with the age of the patient makes it possible to more accurately and indirectly measure the patient's intrapulmonary pressure.
[0142]
Further, if the pressure detection container 85 and the detection branch pipe 84 are prototyped so that the coefficient k in the equation (8) is substantially equal to 1, only a pressure value offset due to the term m is generated. Since the curve changes in the model lung pressure sensor output and the corresponding pressure detection sensor output are the same, it is easy for a doctor to observe and calculation of intrapulmonary pressure and ventilation volume is also facilitated. The waveform data of FIGS. 12 and 14 are substantially the same because the pressure detection container 85 and the detection branch pipe 84 are designed so that the coefficient k is approximately 1.
[0143]
In addition, since the “peak to peak” of the waveform in this case coincides with the “model lung 82” and the “pressure detection container 85”, the ventilation volume of the patient can be simplified indirectly from the previous equation (16). It becomes a management index of artificial respiration of the high-frequency artificial respiration method, and it becomes easy for a doctor to manage treatment of a patient.
[0144]
【The invention's effect】
In the present invention, the inner diameter and length of the detection branch tube and the volume of the pressure detection container are related to the linear function of the insertion side tip pressure of the endotracheal insertion tube and the internal pressure of the pressure detection container during high-frequency artificial respiration. Is set to a value that gives rise to The controller calculates the intrapulmonary pressure of the patient from the detected pressure inside the pressure detection container when performing high-frequency artificial respiration on the patient.
[0145]
Therefore, according to the present invention, it is possible to measure the pressure change and the ventilation volume in the lungs of the patient during the high-frequency ventilation, which is impossible in the past, and it is possible to observe the ventilation state without requiring skill and experience. In addition, since it is not necessary to rely on visual observation as in the prior art, it is possible to grasp a more accurate state, and it is possible to always maintain a more favorable high-frequency artificial respiration.
[0146]
Further, by providing the operation panel for inputting the above-described linear function, it is possible to freely input a linear function determined according to the selected detection branch pipe and pressure detection container.
[0147]
Furthermore, by providing a display unit that displays the calculated intrapulmonary pressure, the intrapulmonary pressure during high-frequency artificial respiration can be known easily and quickly.
[0148]
Similarly, by providing the ventilation amount calculation unit, an external processing device is not required, and the ventilation amount can be grasped only by the configuration of the present invention.
[0149]
Since this invention is comprised and functions as mentioned above, according to this, the outstanding high frequency ventilator which is not in the past can be provided.
[Brief description of the drawings]
FIG. 1 is a block diagram showing a configuration of a high-frequency ventilator according to the present embodiment, showing a state during artificial respiration.
FIG. 2 is a block diagram showing a configuration of a high-frequency ventilator according to the present embodiment, showing a state of a preparatory work stage before performing artificial respiration.
FIG. 3 is a cross-sectional view showing details of the flow control valve disclosed in FIGS.
FIG. 4 is a chart showing an example of dimensions of an endotracheal tube according to age.
FIG. 5 is a block diagram showing a control system of a high-frequency ventilator.
FIG. 6 is a diagram showing a relationship of “internal pressure−ventilation amount” at the time of inhalation supply and at the time of exhalation discharge.
FIG. 7 is an explanatory diagram showing a model of a human lung.
FIG. 8 is an explanatory diagram showing an “airway branching configuration”.
FIG. 9 is a diagram showing the relationship between the dimension of branching and the sum of airway cross-sectional areas for each dimension.
FIG. 10 is an explanatory diagram showing a relationship between branch dimensions and airway resistance by trachea in each dimension.
FIG. 11 is a diagram showing a waveform example of a typical model lung internal pressure when performing high-frequency artificial respiration according to the present embodiment.
12 is a diagram showing the detected pressures of the first and corresponding pressure detection sensors when the oscillating air pressure is 5 Hz in the high frequency ventilator shown in FIG. 1. FIG.
13 is a chart showing test results when the vibration air pressure is 5 Hz in the high frequency ventilator shown in FIG. 1. FIG.
14 is a diagram showing detected pressures of the first and corresponding pressure detection sensors when the oscillating air pressure is set to a frequency of 15 Hz in the high-frequency ventilator shown in FIG. 1. FIG.
15 is a chart showing test results when the vibration air pressure is set to a frequency of 15 Hz in the high-frequency ventilator shown in FIG. 1. FIG.
16 is a chart showing test results when each tube and each sealed container of the high-frequency ventilator shown in FIG. 1 are set for a child, and the vibration air pressure is 5 Hz. .
FIG. 17 is a chart showing test results when each tube and each sealed container of the high-frequency ventilator shown in FIG. 1 are set for a child and the vibration air pressure is set to a frequency of 15 Hz. .
FIG. 18 is a block diagram showing a configuration of a conventional high-frequency ventilator.
[Explanation of symbols]
12 High frequency ventilator
40 controller
42 Memory unit
43 Control panel
44 display
47 Lung pressure calculator
48 Ventilation volume calculator
50 Vibrating air pressure biasing part
60 Patient route
62 Intake air inlet
70 Discharge route
81 Endotracheal tube
82 model lungs
83 Model lung pressure sensor
84 Branch pipe for detection
85 Pressure detection container
86 Compatible pressure detection sensor
X patient

Claims (5)

An inhalation intake section for supplying inhalation containing oxygen to the patient; a vibration air pressure biasing section for biasing the vibration air pressure with a period higher than the breathing cycle of the patient to the inspiration; A high-frequency respirator comprising a patient-side route for guiding the patient to the patient, and an exhaust route for discharging exhaled air containing carbon dioxide from the patient into the atmosphere.
Equipped with an endotracheal insertion tube inserted from the patient's mouth to near the first branch of the trachea at the patient side end of the patient side path;
A branch pipe for detection branched from a midway portion of the discharge path from the patient-side path, a pressure detection container communicating with any of the paths via the branch pipe for detection, and an inside of the pressure detection container A corresponding pressure detection sensor for detecting pressure,
Regarding the inner diameter and length of the detection branch tube and the volume of the pressure detection container, the relationship between the pressure at the insertion end of the insertion tube in the trachea and the internal pressure of the pressure detection container during a high-frequency artificial respiration is a linear function. Is set to a value that produces
And a controller having a function of storing the relationship of the linear function and calculating the pressure at the distal end portion on the insertion side of the endotracheal insertion tube based on the detected pressure from the corresponding pressure detection sensor. Ventilator. The said controller is provided with the memory | storage part which memorize | stores the said linear function, and the intrapulmonary pressure calculation part which calculates the pressure of the insertion side front-end | tip part pressure of the said endotracheal insertion tube with reference to this linear function. Item 2. The high-frequency ventilator according to Item 1. The high-frequency respirator according to claim 1 or 2, wherein an operation panel for inputting the linear function from the outside to the storage unit is provided in the controller. The high-frequency ventilator according to claim 1, 2 or 3, further comprising a display unit for displaying the pressure of the distal end portion of the insertion tube of the endotracheal tube calculated by the controller to the outside. . The said controller is provided with the ventilation volume calculation part which calculates the ventilation volume with respect to the said patient's lungs based on the pressure of the insertion side front-end | tip part pressure of the calculated endotracheal insertion tube. Or the high-frequency respirator of 4.

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