JP4361136B2 - Tissue heat excision system and method using porous electrode structure - Google Patents

Description

表−１は、イオンコントラスト媒体が、同等な抵抗結果を達成し且つ所望の損傷を創り出すのに必要とされる電力を低減できることを示している。
親水性又は疎水性のいずれかの多孔質素材に対して、本システム１０は、本体と組織との中間面に近接したインピーダンスを検出する装置を備えることができる。図９が示すように、インピーダンスは、散布率の増大にもかかわらずインピーダンスが安定する限度点に到達するまで、液体散布の流量割合の増大と共に低減する。インピーダンスを検出することで、最小流量割合Ｒ_MIN（そこでは、インピーダンスが余りにも高い）と最大流量割合Ｒ_MAX（それ以上では起こり得る塩分又は水分の過負荷状態が存在するようになる）との間で散布流量を制御することができる。
本体内の導電性媒体に浸漬された電極３０の表面積は、イオン輸送を高めるために増大される。しかし、本体の小さな形状萎み性や全体的な可撓性等の所望の特徴は、電極寸法に実際的な拘束を課す。
本体２２の細孔４４へ電極３０を接近させると、導電媒体を通したイオン輸送の効率も高める。また、可撓で小さな萎んだ形状を示す構造上の特徴は、これを考慮する上での実際上の制約となる。
本体２２の形成
拡張可能で−萎むことができる本体２２は、ガラス製鋳型の外部又は内部の周りで形成される。この構成では、鋳型の外部寸法は、拡張可能で−萎むことができる本体２２の所望の拡張した形態にマッチしている。鋳型は、所望の壁厚が達成されるまで本体素材の溶液中に所定の手順で浸漬される。鋳型は、次いで成形された拡張可能で−萎むことができる本体２２を残して食刻除去される。
代わりに、拡張可能で−萎むことができる本体２２は、更に押し出し成形されたチューブからブロー成形される。この構成では、本体２２は、接着剤又は熱溶着を使って一端でシール結合される。本体２２の他方端は、開放したままになっている。拡張可能で−萎むことができる本体２２は、鋳型の内側に設置される。高圧ガス又は流体等の膨張本体が、開放したチューブ端から導入される。チューブ状本体２２が鋳型形状を取るように膨らませられると熱に晒される。成形された拡張可能で−萎むことができる本体２２は、次いで鋳型から引き抜かれる。
本体２２の有孔率は、炭酸ガスレーザ、エキシマレーザ、ＹＡＧレーザ、高出力ＹＡＧレーザ、電子粒ボンバードメント等によって成形の前後のいずれかで与えられる。
以前に説明したように、組織切除のための本体２２の電気特徴を改善するために表面の親水性をより高めるように、コーティングや表面処理も施工される。
市場で入手できる多孔質素材も本体２２へと成形される。再生セルロースのような化学工程によって形成される接着性の悪いその種の素材に対して、素材は、浸漬加工（図３２Ａに概略示されており、後で説明する）や、射出成形によって、又は押出し成形中に直径や外形を変えることによって化学的に３次元の形状に成形される。
熱接着やレーザ溶接、超音波溶接及び接着剤による接合が行われるそれら素材に対しては、その素材が使われているシートから３次元の形状を成形するために、これらの接着や溶接の技術を使ういろいろな方法がある。据付け物とマンドレルが、熱と圧力とを関連させて本体２２を成形するのに使用される。
図１９から２３は、多孔質素材２００を切除用本体２２の所望の３次元の形状に成形する上で好ましい方法を示している。図１９に示すように、シート２００が据付け物２０４上の成形腔所２０２に設置される。成形腔所２０２の形状は、本体２２の末端部に対して所望される形状に対応している。図示の実施例では、その形状はほぼ半球状となっている。
図２０が示すように、成形マンドレル２０６はシート２００の一部分２０８を成形腔所２０２内に押し込む。成形マンドレル２０６の形状は、成形腔所２０２の半球形状にマッチしている。マンドレル２０６は、素材部分２０８を間に挟みながら腔所２０２内にしっくりと嵌合する。これで圧力によって所望の半球形状を素材部分２０８にセットする。成形マンドレル２０６か成形腔所２０２のいずれか、又は両方が加熱されて、成形腔所２０２内において素材部分２０８に付加的な熱セットを与えることができる。成形腔所２０２内における圧力と、オプションでの熱は、その素材部分２０８を平坦形状から所望の半球形状に成形する（図２１を見よ）。
先に成形された部分２０８を備えたシートは、据付け物２０４から取外され、仕上げ据付け物２１０（図２２を見よ）上に乗せられる。仕上げ据付け物２１０は、先に成形された部分２０８の形状にマッチした形状を有した末端部２１２を有している。部分２０８は、末端の据付け物端部２１２上にフィットする。
仕上げ据付け物２１０は、図示の実施例でも半球状になっているが、本体２２の近位端部に所望されている形状を有した近位端部２１４を備えている。シート２００は、据付け物２１０の近位端部２１４の周りでうまくひだを造ってまとわり着いている。
仕上げ据付け物２１０は、周りにおいてシート２００の残りの素材がオーバラップしたプリーツ２１８で集合される基部領域２１６を有している。シート２００は、それによって据付け物２１０の全形状にしっかりと沿うことに成る。
仕上げ据付け物２１０は、本体２２の所望の形状を成したシート２００に付加的な熱セットを与える手助けをするために加熱される。成形加工を促進するためにクラムシェルの型（図示されていない）が、更に据付け物２１０の周りに固定される。
さて多孔本体２２（図２３を見よ）として成形されたシート素材は、据付け物２１０から滑らして外される。据付け物２１０の基部領域２１６の周りに集められた素材のプリーツ２１８は、例えば熱接着や超音波溶接によって互いに接着される。これで、直径の絞られた首部分２２０を本体２２に形成し、カテーテルチューブの末端への本体２２の取付けを容易にする。
プリーツ加工の前に、シート端部２１７は、プリーツ加工中に集まって来る素材量を最小に抑制するために部分部分に切り込まれる。プリーツ加工後は、過剰な素材は、折り返されて首部分２２０に接着され、及び／若しくは、さもなければ、首部分２２０と末端部２０８との間に滑らかな変遷部を形成するように切り取られる。
代わりに、据付け物２０４から取り外した後は、先に成形された部分２０８を備えたシート２００は、拡張可能な据付け物２２２（図２４を見よ）の周りに掛け渡される。シート２００の近位端部２１７は、締着部材２２３によって据付け物２２２の首部分の周りにぴったりと締着される。
据付け物２２２は、本体２２に必要とされる形状に気体又は液体を使って拡張されるバルーン（例えば、テフロン材から造られた）等から構成されている。それによってシート２００は、所望の形状を取るように拡張する据付け物２２２によって成形される。
据付け物２２２の拡張前又は拡張中に、熱がシート２００の端部２１７に加えられ、素材を柔軟にして成形加工を補助する。外部圧力もシート２００の近位端部に加えられ、所望の縮径を成した首部分２２０を造るのを補助する。これによって、更に近位端部２１７で素材が『束になる』のを防ぐ助けをする。
据付け物２２２自身は、更に据付け物を拡張するために加熱された気体や液体を使用することによって加熱される。熱は、本体２２の所望の形状にシート２００を付加的に熱セットする。外側のクラムシェルの型（図示されていない）も、成形加工を容易にするために据付け物２２２の周りに固定される。代わりに、食刻で外されるガラス等の素材から成る外側シェルは、所望の最終形状を与えるために使用される。
据付け物２０４又は２２２のいずれかを使用して本体２２の成形に使われるいずれの加熱加工に対しても、細孔寸法が加熱加工中に大幅に変化しないようにヒートシンク（図示されていない）が、先に成形された末端部分２０８を冷やすために使用される。
代わりに、細孔寸法への加熱の影響は、本体２２への成形以前に第１例におけるシート２００の成形に際して推測され、考慮される。例えば、もし細孔が成形中に開くようであれば、細孔は成形中の寸法増大を考慮に入れるべく製造中に比例的により小さく形成される。かくして、所望の細孔寸法は、シートを本体２２に成形している間に最終的に達成されることになる。
成形加工後に、据付け物２２２は収縮されて、引き抜かれる（図２５を見よ）。成形された本体２２２が残る。
なおもう一つ別の代替加工（図２６と２７を見よ）では、本体２２は、２つの予じめ成形された部分２２５を周囲のシーム２２４に沿って結合することによって形成される。図示の実施例では、部分２２５は、その周辺周りの余った素材を切除して、図１９から２１に示されたように半球として成形される。部分２２５は、素材の特徴に基づいて、型成形加工によって同様に予じめ成形されることになる。
２つの部分２２５を結合するシーム２２４は、素材の性質に基づいて熱接着、超音波溶接、レーザ溶接、接着剤接合、縫合等によって形成される。採用される接着や縫合の方法は、確実にシームが気密で液密な部分を形成するように選択される。シーム２２４の引っ張り強度も、多孔質素材の泡立ち点の値を越すべきである。
代わりに、シートから所定寸法に切断された多孔質素材の２つのほぼ円形状の平坦部分が、先行成形を行わずにシームによってそれらの周辺周りで結合される。これで、開放した内部を取囲んだ通常萎んでいるディスクを創り出す。使用中、開放した内部に空気が液体を導入することで、ディスクを本体２２に必要とされる形状に拡張させる。ディスクも、ディスクを所望の形状にさせる内部支持構造体５４（図５から７に概略化されている）を取り囲んでいる。
好ましくは、シーム２２４で半球状の又は平坦な部分２２５を結合した後、シーム２２４に対して越えて広がった余った素材は切除される。尚、組織とシーム２２４の幾分粗くなった面部分との接触は外傷を発生させるので、結合された部分２２５は、好ましくは裏返しされる。裏返しで、シーム２２４を（図２８Ｂが示しているように）本体２２の内部に設けることになり、組織との直接接触から遠ざける。
図２８Ａが示すように、結合されけた部分２２５は、一端２５２に小さな孔２５０をあけて、引っ張りワイヤ２５４を挿入し、それを他方端２５６に取付け、その他方端２５６をその孔２５０を通して引っ張ることによって裏返される。これで、取付けられた半球状部分２２５を内側を外に返すことになる。
上記実施例では、周囲のシーム２２４は、本体２２の軸線周りに延びている。代わりに、（図２９Ａと２９Ｂが示しているように）シーム２２６は、本体２２の軸線に沿って延び、平坦な又は３次元の形に予じめ成形されたいずれかの２つ以上の部分２２８を本体２２に結合することができる。各々が本体部分２２８を支持する嵌合据付け物（図示されていない）が使用され、熱や超音波エネルギーがシーム２２６を造るために加えられている間は部分２２８を静止させておく。
図３０が示しているように、他の軸方向に延びたシーム２３０も、シートをもう一つ別のシートに結合するのではなく、むしろシートを区画するために多孔質素材シート内に設けられる。更に、区画された多孔電極についての詳細については後で説明する。図解の目的で、図３０は、シーム２２６、２３０に沿った区画部分の半球状の突出を幾分誇張して示している。
好ましくは、結果的に得られた本体２２は丁度説明したように裏返され、軸方向に延びたシーム２２６又は２３０を本体２２の内側に設けるようにしている。
図２６から３０に示された部分２２５又は２２８は、平坦な又は３次元形状に予じめ成形されているかを問わず、必ずしも同じ素材から造られる必要が無いことを察知すべきである。異なった多孔特徴の素材は、丁度説明したようにシームによって結合される。代わりに、多孔質素材は、導電性又は電気絶縁性のいずれかである孔の無い素材にシームによって結合される。更に代替構成として、導電素材は、シームによって絶縁素材に結合され、（導電性の）一方が組織に接触し、（電気絶縁性の）他方が血液溜りに露出する二重側部を有した電極本体を提供することができる。事実上、電極本体に関連して使用に適したどんな可撓材も本発明のこの局面に係るシームを使って結合される。更に、複合電極本体を形成するためにシームによって共に結合される部分の数は変えられることも理解すべきである。
勿論、色々な具体的な形状も選択される。好適な形状は、基本的には図２が示しているように末端部が球状輪郭を成した球形状で対称形となっている。しかし、非対称または非球状の形状も使用される。例えば、拡張可能で−萎むことができる本体２２は、カテーテルチューブ１２を取付けるために内側に徐々に湾曲したり、首をつくる平坦な末端輪郭で形成される。後で説明するように、図３３と３４Ｂが示しているように細長い円筒形状も使用される。
図１０は、代替の拡張可能で−萎むことができる多孔本体５０を示している。この実施例では、本体５０は通常拡張された形態の形状を取るように成形された開放気孔の発泡体から構成されている。電極３０は、発泡本体５０内に取り囲まれている。高浸透圧の液体媒体３８が開放気孔を満たしながら発泡本体５０内に導入され、既に説明したように切除エネルギーの所望のイオン輸送ができるようにしている。発泡本体５０を使用したイオン輸送も、もし本体５０が散布率を制御するために発泡本体５０の有孔率よりも小さい有孔率を与えてくれる外部多孔スキン５１（図１０の右側が示しているように）を備えていれば、実行されることに成る。
この構成では、（前に説明したように）摺動シースがカテーテルチューブ１２に沿って前進され、発泡本体５０を萎んだ形態に圧縮する。同様に、シースを引き込めると、圧縮力を取り除く。シースの無くなった発泡本体５０は、跳ねて開き、拡張可能で−萎むことができる本体５０を拡張した形態に戻す。
図示された好適な実施例では、末端の操縦機構５２（図１を見よ）は、展開中でも展開後でも多孔質電極構造体２０の操作を高める。操縦機構５２は変更できる。図示の実施例（図１を見よ）では、操縦機構５２は、ハンドル１８によって保持された外部操縦レバー５８に連結された回転カム輪５６を有している。回転輪５６は、左右の操縦ワイヤ６０の近位端を保持している。ワイヤ６０は、切除エネルギー信号ワイヤ３２と共にカテーテルチューブ１２を通って、末端のチューブ端１６に隣接した弾性可撓ワイヤ又はリーフスプリング（図示されていない）の左側と右側に接続している。更に、この詳細と他のタイプの操縦機構が、参考までにこの明細書に組み込まれているランキスト氏とシンプソン氏の米国特許第５，２５４，０８８号に示されている。
図１に示されているように、操縦機構５２のリーフスプリングは、カテーテルチューブ１２の末端１６内に支持されている。図１が示しているように、操縦レバー５８を前方に動かすと、一方の操縦ワイヤ６０を引いてリーフスプリングを撓めたり湾曲させ、それと共に末端のカテーテル端１６と電極構造体２０とを一方向に撓めたり湾曲させる。操縦レバー５８を後方に動かすと、他方の操縦ワイヤ６０を引いてリーフスプリング６２を撓めたり湾曲させ、それと共に末端のカテーテル端１６と電極構造体２０とを他方向に撓めたり湾曲させる代わりに、図３２Ｃが示すように、操縦可能なリーフスプリング２６８は、多孔本体２２の末端にそれ自から取り付けられている末端取付け具２７０の一部分である。この構成では、リーフスプリング２６８は、多孔本体２２の内側のチューブ２７２内で末端のカテーテル端を越えて延びている。リーフスプリング２６８に取り付けられた操縦ワイヤ６０、６２も、チューブ２７２を通っている。リーフスプリング２６８の近位端は、末端のカテーテル端１６に取り付けられたハブ２７４に取り付けられている。
この構成では、ハンドル１８上の操縦レバー５８の前後移動で、本体２２内においてリーフスプリング２６８を両方向に湾曲させる。リーフスプリング２６８は、末端取付け具２７０を動かし、リーフスプリング２６８が湾曲する方向に多孔本体２２を変形させる。
いずれの構成でも、操縦機構５４は、拡張可能で−萎むことができる本体がその萎んだ形態にあっても、その拡張した形態にあっても使用可能である。
図３２Ａ、３２Ｂは、末端取付け具２７０とリーフスプリング２６８とを多孔本体２２に取り付ける好適な方法を示している。図３２Ａでは、多孔本体２２は、所望の形状を有した拡張可能な据付け物２７８を再生セルロース２７８の溶液内に浸漬することで形成される。そのような拡張可能な据付け物２７８の詳細は、もう一つ別の文脈で既に説明し、図２４、２５に示した。多孔本体２２は、既に説明したように、色々な他の方法で形成されることを察知するべきである。
図３２Ｂが示しているように、据付け物は、近位の首部分２８０と末端の首部分２８２とを有し浸漬形成された多孔本体２２を形成する。本体２２の成形後は、拡張可能な据付け物２７６は、図３２Ｂも示しているように、萎まされて引き抜かれる。
図３２Ｂが示しているように、末端の首部分２８２は、例えば接着剤を使って、又は接着剤接合や熱接合、機械的接合、ネジ、巻き付け、又はこれらのいずれかの組合わせによって取り付けられるスリーブ２８８を使って、末端の取付け具２７０の周りに取付けられる。
取付け具２７０は、既に説明したように、それに予じめ取り付けられたリーフスプリング２６８と、関連した構成要素とを有している。最初取付け具２７０に取り付けられると、本体２２の近位の首部分２８０は、リーフスプリング２６８の反対側の方向に向けられる。
末端の首部分２８２を取付け具２７０に取り付けた後、本体２２は、図３２Ｃが示すようにリーフスプリング２６８を覆うように末端の取付け具２７０の周りで裏返しされる。リーフスプリング２６８の近位端は、末端のカテーテル端１６によって支持されたハブ２７４に取付けられている。裏返された近位の首部分２８０は、次いでスリーブ２８６を使用して末端のカテーテル端に取付けられる。スリーブ２８６は、接着剤接合や熱接合、機械的接合、ネジ、巻き付け、又はこれらのいずれかの組合わせを含んだ色々な方法でカテーテルチューブの周りに取り付けられる。
多孔電極本体をカテーテルの末端に取り付ける色々な代替方法は、『チューブと電線を拡張可能で−萎むことができる構造体に取付ける幹部エレメント』（弁理士ラベル２４５６Ａ−６）の名称の同時係属特許出願に開示されている。
後でより詳細に説明するように、末端の取付け具２７０は、多孔本体２２上の孔の無い導電部分としての働きもできる。同様な取付け具２７０は、同じ目的のために多孔本体２２上の他の場所に配置される。
探針（図示されていない）も、リーフスプリング２６８の代わりに又は組合わせて末端の取付け具２７０に取り付けられる。そこから、探針は本体２２（裏返しに続く）の内側において、カテーテルチューブ１２を通ってハンドル１８上の適当な押し−引き制御器（図示されていない）まで延びている。探針はカテーテルチューブ１２の軸線に沿って移動可能となっており、末端の取付け具２７０上で軸方向に押し引きされ、それによって本体２２を伸ばしたり、又は短縮する。
更に、電極本体への末端の取付け具の取り付けに関する詳細は、『末端で操縦又は操作する拡張可能で−萎むことができる電極構造体』（弁理士ラベル２４５８Ａ−４）の名称の同時係属特許出願に示されている。
図３３と３３Ａ／３３Ｂは、図示のように末端の取付け具に組み合わされ且つ図３２Ｂと３２Ｃに示されているように、末端のカテーテル端１６に取り付けられる細長くて円筒状の例示的な電極本体を示している。
図３３では、本体２９０は、末端と近位の首部分２８０と２８２を形成するために、半径を変えて押出しや浸漬や型成形によって細長い形状に成形される。適当な末端の取付け具２７０（仮想線で示されている）が、末端の首部分２８２内に取付けられており、細長い本体２２は、図３２Ｂと３２Ｃに示されれているように裏返されて組立体を完成する。近位の首部分２８０は、次いで図３２Ｃに示されているように末端のカテーテル端１６に取り付けられている。
図３４Ａと３４Ｂでは、本体２２は、（図３４Ｂに示されているように）一様な半径で押出し成形や型成形や浸漬によって形成された素材のチューブ２９２から形成されている。この構成（図３４Ｂを見よ）では、以前に開示されているように形成されたシーム２９４は、チューブ２９２の末端を閉じている。チューブ２９２の近位端は、末端のカテーテル端１６に取付けるために管状孔２９６の周りでシールされる。代わりに、チューブ２９２の末端は、（図３４Ｂで仮想線で示されている）末端の取付け具２７０の周りでシールされる。後者の場合、チューブ２９２は、カテーテルの末端１６に取り付ける前に末端の取付け具２７０の周りで裏返しにされる。
本体の多孔領域を限定する細孔４４のパターンは変えられる。好ましくは、図２と３に概略示されているように、拡張可能で−萎むことができる本体２２の少なくとも近位側の１／３の表面部分は、細孔４４が無い状態となっている。
拡張可能で−萎むことができる本体２２の少なくとも近位側の１／３の表面部分上に細孔４４を設けていないのは、幾つかの理由で望ましい。この部分は、通常組織とは接触しておらず、その結果、事実上の電極境界が存在していても、何んの目的も果たしていない。更に、この部分も一番小さな直径を示している。もし電気的に導電するものであれば、この部分は、望ましくない最大の電流密度を有することになる。通常組織接触を行っていない最小直径の近位部分を細孔４４の無い状態にしておくことで、組織接触をすることになる拡張可能で−萎むことができる本体２２の末端部分において又はその近くで最大の電流密度が確実に分配されるようにする。
本体２２の末端部分が組織と端部接触するように向けられて切除が行われるものと予測される場合、多孔領域は、勿論拡張可能で−萎むことができる本体２２の末端部の先端の周りに向けられるべきである。この端部に臨む向き合わせのために、多孔領域は、図２と３が示しているように、本体２２の末端部１／３から１／２に被せられ連続したキャップから構成される。しかし、組織との末端部接触が企図されている場合、好適な実施例（図１１を見よ）では、導電多孔領域は本体２２の末端部先端周りの同心状の『雄牛の目』パターンで配列された別々のエネルギー伝達ゾーン６２に区画される。
本体２２の側部領域を組織と接触するように向きを合わせして切除が行われるものと予測される場合、多孔領域は、本体の末端部１／３から１／２の周りで周方向に隔設された軸方向の細長いエネルギー伝達ゾーン６２（図１２を見よ）に好ましくは区画される。
多孔領域が本体２２上の区画されたゾーン６２から構成されると、内部でグループを成したシールされた気のう６４（図１３を見よ）が個別に液体３８を各多孔領域区画部６２に搬送する為に使用される。各気のう６４は、個別に管腔６６と連通しており、それが担当している一つの多孔区画部６２の導電液を受け取る。多数の気のう６４は、更に、幾つかの、しかし全部の気のう６４ではないが液体で選択的に膨らませることで、膨らんだ本体２２の形状をより具体的に制御する能力も与えてくれるものである。
気のう６４は、別々に成形されて本体２２内に挿入されるか、又はそれらは、拡張可能で−萎むことができる本体２２を型成形している間に一体的に形成される。
図１２が示しているように、区画された多孔ゾーン６２は、更に拡張可能で−萎むことができる本体２２の萎みに関連して使用するのにも良く適合している。この構成では、無孔電極は、ひだのある又は祈り目のある領域６８から構成されている。これらの領域６８を創り出すために、本体２２の型は、拡張可能で萎むことができる素材が所望の領域６８に沿って若干より薄く形成されたり、凹み付けされたり、リブ付けされるように予じめ成形された表面形状を有している。拡張可能で−萎むことができる本体２２は、一貫して一様に本体２２を自からの上に周方向で萎ませながら、これらのひだのある領域６８の周りで萎む。かくして、結果的に生じる萎んだ形状は、より一様で且つコンパクトにされる。
図１２に示されている萎み可能な本体２２も、多孔領域の他のパターンにも使用されることを察知すべきである。ひだのある領域６８には、更にもし必要ならば細孔も設けられる。
図１４は、二重の作用を果たす拡張可能で萎むことができる電極構造体７０の実施例を示している。構造体７０は、前に説明したように内部電極３０を収容した拡張可能で萎むことができる本体２２を有している。本体２２は、導電流体３８を収容しており、また更に丁度説明したように電気エネルギーのイオン輸送を可能にする一つ以上の多孔領域６２も有している。
図１４に示されている構造体７０は、更に本体２２の表面上に一つ以上の導電領域７２も有している。（図１４が示しているような）一実施例では、孔の無い導電領域７２は、スパッターリングや、蒸着や、イオンビーム被覆や、沈着された種層を覆う電子メッキや、又はこれらの加工の組合わせによって拡張可能で−萎むことができる本体２２上に被覆される金、プラチナ、プラチナ／イリジウムなどの金属から構成される。代わりに、孔の無い導電領域７２は、本体の表面に固定された薄い箔からも構成される。更に代わりに、孔の無い導電領域７２は、１箇所以上で多孔本体２２によって支持された（図３２Ｃに示された末端の取付け具のような）硬質な取付け物からも構成される。本体２２内の信号ワイヤ（図示されていない）は、無孔電極に電気的に接続されている。信号ワイヤは、ハンドル１８によって保持されたコネクター３８に接続するためにカテーテルチューブ１２を通っている。
好適な実施例（図１５を見よ）では、孔の無い導電領域７２は本体の内部に通されて電気接続部の所望の点で、本体２２を貫いて蛇行された絶縁信号ワイヤ２６から構成されている。蛇行貫通されたワイヤ２６の末端部の電気絶縁は除去され、導電領域７２としての働きをするように好ましくは平坦化される電線部分を露出する。平坦化された領域では、導電接着剤７３によって本体２２に固定取り付けされる。接着剤７３は、更に好ましくはワイヤ２６が通っている本体２２の領域にも塗られてそれをシールする。同じ信号ワイヤ２６は、多数回に渡って本体２２を貫通して蛇行されて、必要に応じて多数の領域７２を形成する。
無孔電極７２と、関連した信号ワイヤとを拡張可能で−萎むことができる電極本体２２に取り付ける色々な方法については、『電極構造体の向上された電気接続』（弁理士ラベル２４５８Ａ−５）の名称の同時係属特許出願に説明されている。
無孔電極７２は、心筋組織における電気的活動を検出するのに使用される。検出された電気的活動は、外科医による分析のための電位を処理する外部制御器に送られる。その処理によって、潜在的な不整脈の病巣を特定するために電位又は消極事象のマップが作られる。無孔電極７２で一旦特定されると、多孔領域６２は、病巣を切除するために前に説明したように無線周波数エネルギーを送るために使用される。
代わりに、又は電気的活動の検出との組合わせで、無孔電極７２は、ペース信号を送るために使用される。このように、無孔電極はペースマッピング又は同調化マッピングを実施できる。
好ましくは（図１６を見よ）、本体２２の表面は、単極又は双極の検出に又はペース取りに適した電極１００を保持している。これらの電極１００は、本体２２の内側表面上に配置されているが、検出や又はペース取りのためのそれらの作用は、本体２２の良好な導電特徴の理由で害されることはない。
異なった電極の設置が、単極又は双極の検出に又はペース取りのために使用される。例えば、対になった２−ｍｍの長さで１−ｍｍの幅の電極が本体２２の内側表面上に設けられる。接続ワイヤ１０２がこれらの電極１００に取り付けられる。高浸透圧の溶液がこれらの電極を短絡させるのを防ぐために、それには、（例えばエポキシや接着剤等の）電気絶縁材１０４で覆われなければならない。好ましくは、電極間の距離は約１ｍｍで、確実に良好な双極の活動電位図が得られるようにしている。これら内部電極の好ましい設置箇所は、構造体２２の末端部の先端と中心とである。更に、多数のゾーンが使用される場合は、切除領域間に電極を設けるのが望ましい。
外科医が本装置を螢光透視法の下で目標とする部位に案内できるように不透明なマーカー１０６を本体２２の内側表面上に設けるのも望ましい。この目的には、いずれの高原子重量材も適している。例えば、プラチナやプラチナ／イリジウムがマーカー１０６を作るのに使用される。これらのマーカー１０６の好ましい設置箇所は、構造体２２の末端部の先端と中心である。
それによって、図１４に示されている拡張可能で−萎むことができる構造体７０は、『固体の』無孔電極７２の使用を『液体の』又は多孔な電極６２と組合わせている。拡張可能で−萎むことができる構造体は、一つの電極作用を使う治療目的のために心筋組織のマッピングと、異なった電極作用を使う治療目的のために心筋組織の切除とを実施可能にする。
代替実施例では、構造体７０の無孔電極７２は、無線周波数エネルギーを送って組織を切除するために多孔領域６２と共に直列状態で使用される。この構成では、領域７２を受け持つ信号ワイヤは、発生器４０に電気的に接続されており、一つ以上の領域７２による伝達のために無線周波数エネルギーを搬送する。同時に、内部電極３０は、多孔本体を通る媒体３８による伝達のために無線周波数エネルギーを受け取る。領域７２を取囲んでいる多孔構造体を横切ってイオン輸送を行うことで、切除電極の有効表面面積を拡げる。
この実施例では、図１４に示されている拡張可能で−萎むことができる構造体７０は、それで検出とマッピングのために第１の有効表面面積を有した電極７２の使用を組合わせている。第１の有効表面面積は、電極７２を取り囲んだ多孔構造体を横切って高浸透圧液体のイオン輸送を行うことによって切除を行う切除目的の為に選択的に増大される。
もし液体散布が細孔を通って行われれば、内部電極３０は、それら領域の有効電極表面面積を増大するように要求されることはない。それら領域が無線周波数エネルギーを伝達する時に行われる細孔を通したイオン媒体の液体散布は、それ自身それら領域７２の有効な伝達表面面積を増大するのに十分な働きをする。しかし、もしイオン輸送が実質的に液体散布無しで行われれば、無線周波数が伝達のために外部領域７２に供給されている時と同時に内部電極３０を使用して無線周波数エネルギーを伝達することも有利になると思われる。
この実施例では、領域７２はそれら自身多孔導電材から造られることも知るべきである。このように、イオン輸送は、領域７２自身を横切って行われ得る。
前に説明したように（図１を見よ）、制御器３２は、好ましくは、発生器３０から本体２２内に支持された電極への無線周波数切除エネルギーの搬送を制御している。好適な実施例（図２を見よ）では、多孔電極構造体２０は、制御器３２に接続された１つ以上の温度検出エレメント１０４を保持している。
温度検出エレメント１０４は、サーミスターやサーモカップル、又はその種の形をしている。温度検出エレメント１０４は、電極構造体２０の外部と熱伝導接触しており、切除中に構造体２０の外側の組織の状況を検出する。
温度検出エレメント１０４によって検出された温度は、制御器３２によって処理される。入力された温度に基づいて、制御器３２は電極３０による無線周波数エネルギー伝達の時間と電力レベルとを調節し、所望の損傷パターンと他の切除目的を達成する。
温度検出エレメント１０４を拡張可能で−萎むことができる電極本体に取り入れる色々な方法については、『電極構造体用の向上された電気接続』（弁理士ラベル２４５８Ａ−５）の名称の同時係属特許出願に説明されている。
図３１Ａ、３１Ｂと３１Ｃに示すように、温度検出エレメント１０４は、更にシート２６０、２６２を本体２２に共に結合しているシーム２５８の近くに又はその内部に配置される。そのようなシーム２５８の形成については、既に説明したし、更に図２６から３０に示している。
図３１Ａと３１Ｂに示されているように、各温度検出エレメント１０４は、一方のシート２６０上に設置され、次いで他方のシート２６２によって覆われる。２枚のシート２６０、２６２は、次いで共に結合されて本体２２を形成する。シーム２５８は、検出エレメントを包囲している。各検出エレメント１０４の信号ワイヤ２６４は、図３１Ａと３１Ｂが示しているように、シート２６０、２６２の外部までシーム２５８の無い所を延びている。
前に説明したように、本体２２は好ましくは裏返される（図２８Ａを見よ）。図３１Ｃが示すように、裏返しでシーム２５８と包囲された信号ワイヤ２６４の両方を本体２２の内部に設けることになる。信号ワイヤ２６４は、制御器３２に接続するために首部分２６６に通されている。
温度検出エレメント１０４の代わりに又はそれに加えて、ペース用電極と検出用電極を、図３１Ａから図３１Ｃに示されているように、裏返された電極本体２２のシーム２５８内に包囲することもできる。そのような構成では、裏返し後に、本体素材と組織との間に企図した接触が行われる素材表面に接近したシーム内にそれらが配置されるように電極を位置決めするのが好ましい。
更に、多くの温度検出エレメントを使って制御される多くの切除エネルギーの伝達器の使用についての詳細は、１９９４年８月８日に出願され且つ『多くの温度検出エレメントを使用して組織切除を制御するシステムと方法』の名称の同時係属米国特許出願番号第０８／２８６，９３０号に開示されている。
例−１
生体外での分析
３つの電極構成について分析された：
（１）分子量のカットオフを１２，０００−１４，０００ドルトンとし且つ内部液体媒体として９％の塩水溶液を使用して（スペクトラによって製造された）再生セルロースから造られた透析チューブから構成された１３ｍｍディスク状本体を有した本発明に係る多孔な拡張可能で−萎むことができる電極構造体、
（２）１３ｍｍの直径を有したスパッタリングされたプラチナのディスク状の電極本体と、
（３）１０ｍｍの直径を有したアルミ箔製の拡張可能で−萎むことができる半球状電極構造体である。
高温の温度条件が存在した領域において、動物（羊）の組織内に０．５ｍｍだけサーミスターが埋め込まれた。電極（１）、（２）に対して、最高温度がエッジ加熱効果によってディスク状本体のエッジに発生した。電極（３）に対しては、最高温度が電流密度の最も大きい半球状本体の末端部の先端で発生した。
無線周波数の電磁電力は、サーミスター温度を６０℃、７０℃、８０℃又は９０℃に維持するように調整された。全ての損傷の直径は、組織の脱色をマークする６０℃の等温線に基づいて計測された。
次の表−２は、生体外での分析で観察された結果をリストアップしている。

多孔電極構造体（１）は、電極（２）のような正規の金属切除電極と比較して対流式冷却によって最小限の影響しか受けなかった。これは、特定の損傷発生中に流体流量を変え、且つ多孔電極構造体に関してサーミスター温度を維持するのに何ら電力を変化させる必要が無い事を観察して分かった。
表−２は、組織温度に基づいて電力を調節する場合に、多孔電極構造体が、金属被覆された電極構造体と少なくとも同じくらい大きく損傷を創り出したのを示している。多孔電極構造体は、更に金属被覆された電極構造体と比較してリーズナブルなインピーダンスレベルを有していた。
この例は、多孔電極構造体が心外膜の、心内膜の又は壁内の基質を切除するために制御された状態で１．０ｃｍより深い損傷を創り出すことができることを示している。
多孔電極構造体を形成する透析チューブは、高い吸水特徴を有している。透析チューブは水に晒されると大幅により可撓的になる。分子量のカットオフは、１２，０００から１４，０００ドルトンであった。より大きな又はより小さな分子量のカットオフについてはスペクトラム社から入手できる。テストされる透析チューブに対する分子量カットオフから予測細孔寸法への変換は、『透析／超瀘過』の名称の医療器具小冊子９４／９５（ｐ１０）から得られるように、１００，０００ドルトンは０．０１μｍに等しく、５０，０００ドルトンは０．００４μｍに等しく、１０，０００ドルトンは０．００２５μｍに等しく、５，０００ドルトンは０．００１５μｍに等しくなっている。
透析チューブは、小さな細孔寸法にもかかわらず親水特徴と大きい有孔率を有している。結果的に、泡立ち点の値は非常に高く、また抵抗は無線周波数エネルギーの供給中に液体流を必要としない程度に十分に低いものとなっている。
例−２
有限要素分析
３次元の有限要素モデルが、２８．４ｍｍの全長で、６．４ｍｍの直径で０．１ｍｍの本体壁厚の細長い形状を有した多孔電極構造体に対して創られた。０．２ｍｍの直径の金属ワイヤが、ＲＦエネルギー源に接続された内部電極としての働きをするように本体の長さに沿って内部に延ばされた。本体は、５．０Ω・ｃｍの電気抵抗を有した９％の高浸透圧溶液が充満された。構造体の多孔本体が、電線として作られた。心臓組織とのしっかりとした接触が、電極の下方の平面に横たわっている電極本体の全長に沿って確保された。血液との接触が、電極の上方平面に横たわっている電極本体の全長に渡って確保された。血液領域と組織領域は、各々１５０と５００Ω・ｃｍの抵抗を有していた。
電極本体に対して、１．２ｋ−Ω・ｃｍと１２ｋ−Ω・ｃｍの抵抗に基づいて分析が行われた。
表３は、ＲＦ切除電力が電極の多孔本体に対して色々な電力レベルで、また色々な抵抗レベルで多孔電極に加えられた時の最高組織温度の深さを示している。

図１７は、多孔本体が１．２ｋ−Ω・ｃｍの抵抗を有している場合、電極に５８ワットで２４０秒間電力が加えられた時の温度プロフィールを示している。図１７における５０℃の等温線の深さは１．４ｃｍである。
図１８は、多孔本体が１２ｋ−Ω・ｃｍの抵抗を有している場合、電極に４０ワットで２４０秒間電力が加えられた時の温度プロフィールを示している。図１８における５０℃の等温線の深さは１．０ｃｍである。
全ての場合、最高温度は、組織と細長い多孔構造体の両端エッジとの間の中間面に位置している。これは、多孔本体の細長い形状に対する温度検出エレメントの好適な場所は、本体の各端部エッジである。好ましくは、各エッジは、少なくとも一つの温度検出エレメントを保持すべきであり、多数の温度検出エレメントは、少なくともそれらの一つが組織に確実に面向するように対角状に対向した側に配置されるべきである。
データは、更には、金属面電極で観察されるように最も高温な領域が、組織内に深くは移動され無いことも示している。最も高温な領域は、直接検出のために組織−電極本体の中間面に一貫して存在している。幾分より大きい差が、計装する或る他の局面において遭遇されるのであるが、この特長は、検出された温度と事実上最も高温な組織温度との間の差をできれば理論的に０℃に削減することになる。
なるより高い抵抗を有した本体を備えた多孔な電極本体（図１８を見よ）は、より低い抵抗値（図１７を見よ）を有した多孔な本体に比較して見たところ、より一様な温度プロフィールを発生した。電極本体の抵抗を高めたことで組織−電極本体の中間面で発生された付加的加熱の熱によって、同じ最高温度に到達するのに必要とされた電力はより少ないものであった。結果は、損傷深さが浅くなったことであった。
以前に説明したように、本体２２の抵抗を選択することによって、外科医は、損傷形状に大きな影響を与えることができる。小さな抵抗の本体２２を使用することで、結果的により深い損傷を発生させるが、またその反対も起きる。次の表−４は、経験的なデータに基づいて、本体抵抗と損傷深さとの間の関係を説明している。

多孔電極構造体の削減された熱伝導率によって、無孔金属面電極と比較すると、損傷形成は、電極周りの動的な血液流状況に対して感度が小さいものと予測される。多孔電極を介した切除エネルギーの供給は、特に浅い心房用損傷が要求される時に所望の損傷特徴を得るためにより精密に制御される。
次の表−５は、経験的なデータに基ずいており、血液流量における変化によって対流式冷却条件に対する多孔電極構造体の感度低下を証明している。

多孔な電極構造体を使用することで、構造上の長所が与えられる。それは、金属の導電シェルを拡張可能で−萎むことができる本体の外側に設置することで関連事象として起こり得る付着についての問題点を切り離してくれることである。多孔電極構造体も、外部の導電材への組織の付着が起り得る潜在的な問題も回避してくれる。
これらの構造上の長所に加えて、切除作業の温度制御が改善される。組織を切除するのに在来の金属電極を使うと、組織−電極の中間面は、取囲んでいる血液流によって対流式で冷却される。これらの対流式冷却作用に依って、最高組織温度の領域は、組織においてより深く位置される。結果的に、金属電極エレメントに関連した検出エレメントによって検出された温度条件は、実際の最高組織温度を直接的に反映していない。この状況下では、最高組織温度の条件は、実際の検出された温度から推断され、予測されなければならない。多孔電極構造体２０又は７０を使用して、周囲の血液流によって組織−電極の中間面を対流で冷却するのを最小へと減ずる。結果的に、最高温度の領域は、組織と多孔電極との間の中間面に位置されることになる。結果的に多孔電極エレメントに関連した検出エレメントによって検出された温度条件は、より精密に実際の最高の組織温度を反映することになる。
例−３
生体外の実験が、親水性素材（Ｈｐｈｉ）対疎水性素材（Ｈｐｈｂ）を多孔組織切除エレメントとしてそれらを使った項目で比較するために実施された。表−６は、その結果を要約している。

表−６は、細孔寸法が親水性素材を使用して小さくされ、それによって多孔質素材を通した液体散布を最小限にし、又は停止し、他方依然として膜を通してイオン輸送ができるようにしておることを証明している。
疎水性多孔質素材は、高い抵抗の多孔電極を実現できるようにする。他方、親水性多孔質素材は、低い抵抗の多孔電極を実現できるようにする。
所望の損傷特徴を得る
上記表が証明しているように、同じ拡張可能で−萎むことができる多孔電極構造体２０は、広くて浅い又は大きくて深いいずれかの損傷を選択的に形成できる。色々な方法論が、この結果を得るべく無線周波数エネルギーの供給を制御する為に使用される。
Ａ．Ｄ50C作用
一つの代表実施例では、制御器（コントローラー）２は、外科医から（ｉ）所望の損傷が生存可能な組織と生存不可能な組織との間の境界深さまで組織−電極の中間面の下で広がる範囲、及び／若しくは（ｉｉ）組織−電極の中間面と境界深さの間の損傷内で発生される最高組織温度の項目における所望の治療結果を受け取る入力部３００（図１を見よ）を制御器４２は有している。
制御器４２は、更に、損傷境界深さや切除電力レベル、切除時間、実際の表面下の組織温度及び電極温度の間において観察された関係を相関させる関数を保持する処理エレメント３０２（図１を見よ）を有している。処理エレメント３０２は、所望の治療結果をその関数と比較し、その比較に基ずいて操作条件を選択して、上記の実際の又は予測された表面下の組織温度を越さずに所望の治療結果をを達成する。
処理エレメント３０２によって選択された操作条件は、切除電力レベルの制御や切除時間の選択された目標切除時間への制限、上記最大切除電力レベルを受ける切除電力レベルの制限、及び／若しくは領域４４と組織との間の所望の百分率の接触の予測を含んだ本体２２の多孔領域４４の向き等の切除工程の色々な局面を制御できる。処理エレメント３０２は、実際の最高組織温度を検出する為に組織内に入り込んで拡張可能で−萎むことができる構造体２０によって保持された又は他にはそれに関連した温度センサーに依存している。代わりに、処理エレメント３０２は、操作条件に基ずいて最高組織温度を予測できる。
好適な代表実施例では、電極構造体２０は、領域４４の熱塊の瞬間的な局部温度（Ｔ１）を検出少なくとも一つの温度検出エレメント１０を保持している。或る時間における温度Ｔ１は、発生器４０によって電極３０に供給される電力の関数である。
損傷の特徴は、Ｄ50Cと呼ばれる５０℃の等温領域の組織面下の深さの表現で表現される。深さＤ50Cは、多孔領域４４の物理的特徴（即ち、その電気と熱の伝導率と抵抗と寸法）と；組織と多孔領域４４との間の接触百分率と；領域４４の熱塊の局部温度Ｔ１と；内部電極３０によって伝達されるＲＦ電力の大きさ（Ｐ）と；組織がＲＦ電力に晒される時間（ｔ）の関数でる。
所望の損傷深さＤ50Cに対して、安全について更に考慮するとマトリックスにリストアップされた操作条件中における最適操作条件の選択を拘束することになる。原則的な安全上の拘束条件は、最高組織温度ＴＭＡＸと最高電力レベルＰＭＡＸである。
最高組織温度条件ＴＭＡＸは、深くて広い損傷を与えるのに十分に高いが（一般に約８５℃と９５℃との間）然し安全上約１００℃未満で、組織乾燥や組織微細破裂を起こすことが知られている温度範囲以内に入っていることである。ＴＭＡＸは、中間面とＤ50Cとの間の電極−組織の中間面の下方の或る距離の所で生じることが認識されている。
最高電力レベルＰＭＡＸは、内部電極３０の物理的特徴とＲＦ発生器４０の発電能力とを考慮している。
これらの関係は、上記例が図解しているように、制御された実際のまた模擬実験される条件下で経験的に及び／若しくはコンピュータモデルによって観察される。或る多孔領域４４に対するＤ50C作用は、上記値のマトリックスにリストアップされている全て又は幾つかの表現や、経験的なデータ及び／若しくはコンピュータによるモデル化から引き出されたそれらの関係表現で表現される。
処理エレメント３０２は、上述のようにｔ＝１２０秒とＴＭＡＸ＝９５℃に対してまた一連の他の操作条件に対してＤ50C温度境界関数を定義するこのマトリックスの操作条件をメモリーに有している。
外科医は、更に、上記の識別符号を使用して構造体２０の特徴を識別し、所望の最大ＲＦ電力レベルＰＭＡＸや；所望の時間ｔや；所望の最高組織温度ＴＭＡＸを設定するのに入力部３００も使用する。
これらの入力に基ずいて、処理エレメント３０２は、所望の治療結果をマトリックスに定義されている関数と比較する。発生器４２は、関数の変数を制御して上記ＴＭＡＸを越さずに所望の治療結果を達成するように操作条件を選択する。
それで、この構成は、外科医が事実上所望のＤ50Cを特定することで『ダイアルで損傷』形成できるようにする。
更に、Ｄ50C作用を引き出す詳細と、所望の損傷パターンを得る為のそれの使用とは、ここに参考までに組み込まれている『本体組織の切除中に所望の損傷特徴を得るシステムと方法』の名称で、１９９５年５月１日出願の係属米国出願番号第０８／４３１，７９０号に見ることができる。
Ｂ．区画された領域：デュティサイクル制御
色々なＲＦエネルギー制御方法が、図１１（ゾーンの軸方向に隔設された雄牛の目パターン）と図１２（周囲方向に隔設されたゾーン）とに示されている区画された多孔パターンと関連して使用される。説明目的に、（電極領域とも呼ばれている）４４は、Ｅ（Ｊ）で符号表示されるが、Ｊは或るゾーン４４（Ｊ＝１からＮ）を表示している。
前に説明したように、各電極領域Ｅ（Ｊ）は、Ｊが或るゾーン４４を表し且つＫが各ゾーン上の温度検出エレメントの数（Ｋ＝１からＭ）を表しているＳ（Ｊ、Ｋ）で表示される少なくとも一つの温度検出エレメント１０４を有している。
このモードでは、デュティサイクル１／Ｎの多数のパルスでＲＦ電力を供給するように適当な電力切り替え中間面を介して条件設定される。
パルス化された電力供給で、各個別の電極領域Ｅ（Ｊ）に送られる電力量（Ｐ_E(J)）は次のように表現される。
Ｐ_E(J)αＡＭＰ_E(J) ²×ＤＵＴＹＣＹＣＬＥ_E(J)
その場合：
ＡＭＰ_E(J)は、電極領域Ｅ（Ｊ）に送られるＲＦ電圧の振幅であり、
ＤＵＴＹＣＹＣＬＥ_E(J)は、次のように表現されるパルスのデュティサイクルである：
ＤＵＴＹＣＹＣＬＥ_E(J)＝ＴＯＮ_E(J)／（ＴＯＮ_E(J)＋ＴＯＦＦ_E(J)）
その場合：
ＴＯＮ_E(J)は、各パルス期間中に電極領域Ｅ（Ｊ）がエネルギーを発する時間であり、
ＴＯＦＦ_E(J)は、各パルス期間中に電極領域Ｅ（Ｊ）がエネルギーを発しない時間である。
式（ＴＯＮ_E(J)＋ＴＯＦＦ_E(J)）は、各電極領域Ｅ（Ｊ）に対するパルス期間を表している。
このモードでは、発生器４０は、各電極領域に対して１／Ｎのデュティサイクル（ＤＵＴＹＣＹＣＬＥ_E(J)）集中的に発生できる
発生器４０は、前のパルスのデュティサイクルの終わりが次のパルスのデュティサイクルの始めに若干重なるように隣接電極領域に対して連続した電力パルスを順番に発生する。パルスのデュティサイクルにおけるこの重なりは、なんら中断期間が連続した電極領域間でパルス切り替え中に解放回路によって生じないようにして確実に発生器４０が連続的に電力を供給できるようにするものである。
このモードでは、温度制御器４２は、各電極領域（ＡＭＰ_E(J)）のＲＦ電圧の振幅に対して個々に調節を行い、それ発生器４０によって制御されてデュティサイクル中に各電極領域に送られる切除エネルギーの電力Ｐ_E(J)を個別に切り替える。
このモードでは、発生器４０は、連続したデータ獲得サンプリング期間に循環する。各サンプリング期間中には、発生器４０は、個々のセンサーＳ（Ｊ、Ｋ）と、制御器４２によって出力されるような検出エレメント１０４によって検出された温度符号ＴＥＭＰ（Ｊ）（Ｓ（Ｊ、Ｋ）の内最も高いもの）とを選択する。
一つより多い検出エレメント１０４が或る電極領域に関連している場合（例えばエッジに位置した検出エレメントが使用される場合）、制御器４２は或る電極領域に対して全ての検出された温度を登録し、これらの内からＴＥＭＰ（Ｊ）を構成する最も高い検出温度を選択する。
このモードでは、発生器４０は、各データ獲得期間中に各電極Ｅ（Ｊ）で局部的に検出された温度ＴＥＭＰ（Ｊ）を外科医によって確立された設定点温度ＴＥＭＰSETと比較する。この比較に基ずいて、発生器４０は、その電極領域と全ての他の電極領域に対してＤＵＴＹＣＹＣＬＥ_E(J)を維持している間に、電極領域Ｅ（Ｊ）に供給されたＲＦ電圧の振幅ＡＭＰ_E(J)を変え、設定点温度ＴＥＭＰ_SETにＴＥＭＰ（Ｊ）を発生して維持する。
設定点温度ＴＥＭＰSETは、外科医の判断と経験的なデータに応じて変わり得る。心臓切除の為の代表的な設定点温度ＴＥＭＰは、７０℃が代表的な好ましい値であるが、４０℃から９５℃の範囲にあると思われる。
発生器４０がＡＭＰ_E(J)を制御する方式は、比例制御方法、比例積分導出（ＰＩＤ）制御方法又はあいまい理論制御方法を採用できる。
例えば、比例制御方法を使用して、もし第１検出エレメントによって検出された温度がＴＥＭＰ（１）＞ＴＥＭＰ_SETであれば、発生器３０によって発生された制御信号が第１電極領域Ｅ（１）に加えられるＲＦ電圧の振幅ＡＭＰ_E(1)を低減するが、他方で第１電極領域Ｅ（１）に対するデュティサイクルＤＵＴＹＣＹＣＬＥ_E(1)を同じに維持している。もし第２検出エレメントＴＥＭＰ（２）によって検出された温度がＴＥＭＰ（２）＜ＴＥＭＰ_SETであれば、発生器３０の制御信号は、第２電極領域Ｅ（２）に加えられるパルスの振幅ＡＭＰ_E(2)を増加するが、他方で第２電極領域Ｅ（２）に対するデュティサイクルＤＵＴＹＣＹＣＬＥ_E(2)をデュティサイクルＤＵＴＹＣＹＣＬＥ_E(1)などと同じに維持している。もし或る検出エレメントによって検出された温度が設定点温度ＴＥＭＰ_SETにあれば、関連した電極領域に対してＲＦ電圧の振幅になんら変化を発生させない。
発生器４０は、連続したデータ獲得期間中に電圧差入力を連続的に処理して各電極領域Ｅ（Ｊ）においてＡＭＰ_E(J)を個々に調節するが、他方で集合的なデュティサイクルを全ての電極領域Ｅ（Ｊ）に対して同じに維持している。このように、そのモードは、温度の所望の一様な状態を切除エレメントの長さに渡って維持している。
比例積分微分（ＰＩＤ）制御技術を使用して、発生器は、或るサンプリング期間で生じる瞬間的な変化ばかりではなく前のサンプリング期間で生じた変化とこれらの変化が時間経過に伴って変化する割合も考慮に入れている。かくして、ＰＩＤ制御技術を使用して、前の瞬間的な差に比較して差がより大きく成るのか小さく成るのかに基ずいて、また前のサンプリング期間以来差が変化している割合がより大きく成っているのか小さく成っているのかに基ずいて、発生器は、ＴＥＭＰ（Ｊ）とＴＥＭＰ_SETとの間の或る比例して大きい瞬間的な差に異なって反応する。
更に、温度検出に基ずいて区画された電極領域に対して個別に振幅の／集合的なデュティサイクルを制御する詳細については、ここに参考までに組み込まれている『多数の温度検出エレメントを使用して組織切除を制御するシステムと方法』の名称で、１９９５年５月１２日出願の係属米国出願番号第０８／４３９，８２４号に見ることができる。
Ｃ．区画された領域：差分温度での不能化
この制御モードでは、制御器４２は、各データ獲得位相の終わりにその位相に対して最高（ＴＥＭＰ_SMAX）と成っている検出された温度を選択する。制御器４２は、更に、最低（ＴＥＭＰ_SMIN）と成っている検出された温度をその位相に対して選択する。
発生器は、選択された最も高い検出温度ＴＥＭＰ_SMAXを選択された高い設定点温度ＴＥＭＰ_HISETと比較する。その比較で、比例、ＰＩＤ又はあいまい理論制御方法を使用して、全ての電極領域に対してＲＦ電圧の振幅を集中的に調節する制御信号を発生する。
比例制御実施計画では：
（ｉ）もしＴＥＭＰ_SMAX＞ＴＥＭＰ_HISETであれば、制御信号は全ての領域に送られるＲＦ電圧の振幅を集中的に減ずる。
（ｉｉ）もしＴＥＭＰ_SMAX＜ＴＥＭＰ_HISETであれば、制御信号は全ての領域に送られるＲＦ電圧の振幅を集中的に増大する。
（ｉｉｉ）もしＴＥＭＰ_SMAX＝ＴＥＭＰ_HISETであれば、全ての領域に送られるＲＦ電圧の振幅になんら変化が無い。
発生器は、振幅制御の目的に検出温度ＴＥＭＰ_SMAX、ＴＥＭＰ_SMIN又は間の温度のいずれか一つを選択してこの温度条件を予め選択された温度条件と比較できることを察知すべきである。
発生器は、或る局部温度ＴＥＭＰ（Ｊ）とＴＥＭＰ_SMIN間の差に基ずいて領域への電力供給を制御する。この実行で、局部検出温度ＴＥＭＰ（Ｊ）とＴＥＭＰ_SMIN間の差を演算し、この差を選択された設定点温度差ΔＴＥＭＰSETと比較する。そので、電極領域への電力供給を制御する制御信号を発生する。
もし或る電極領域Ｅ（Ｊ）に対する局部温度ＴＥＭＰ（Ｊ）が、ΔＴＥＭＰ_SETと同じだけ又は（即ち、もしＴＥＭＰ（Ｊ）−ＴＥＭＰ_SMIN≧ΔＴＥＭＰ_SETであれば）ΔＴＥＭＰ_SETより大きく最低の検出された温度ＴＥＭＰ_SMINを越せば、発生器は、或る所定の領域Ｅ（Ｊ）への電力供給を遮断する。ＴＥＭＰ（Ｊ）−ＴＥＭＰ_SMIN＜ΔＴＥＭＰ_SETの場合、発生器はその或る所定の領域Ｅ（Ｊ）への電力供給を復帰させる。
代わりに、ＴＥＭＰ（Ｊ）とＴＥＭＰ_SMINを比較する代わりに、発生器は、ＴＥＭＰ_SMAXとＴＥＭＰ_SMINを比較することができる。ＴＥＭＰ_SMAXとＴＥＭＰ_SMINとの間の差が所定量ΔＴＥＭＰ_SETに等しいか、又は越えている場合、発生器は、ＴＥＭＰ_SMINが存在している領域を除いて全ての領域を遮断する。ＴＥＭＰ_SMAXとＴＥＭＰ_SMINとの間の温度差がΔＴＥＭＰ_SETより小さい場合に、発生器３０はこれらの領域を復帰させる。
更に、差分温度での不能化を使用する詳細については、ここに参考までに組み込まれている『多数の温度検出エレメントを使用して組織切除を制御するシステムと方法』の名称で、１９９４年８月８日出願の係属米国特許出願番号第０８／２８６，９３０号に見ることができる。
Ｄ．ＰＩＤ制御
多孔質電極構造体に関しては、血液溜りによる対流冷却の最小限の作用で、予想された温度に代えて、最高組織温度ＴＭＡＸとして実際の検出された温度条件を使用できるようにしている。これによって、そのような構造体も自ら比例積分導出（ＰＩＤ）制御技術を使用している。これら電極構造体に関連して利用可能な図解のＰＩＤ制御技術は、『監視と制御の為に時間で変わる設定点温度曲線を使用する組織加熱と切除システムと方法』の名称で、１９９４年６月２７日出願の係属米国特許出願番号第０８／２６６，０２３号に開示されている。
最後に、多孔な拡張可能で−萎むことができる本体内部に配置された内部電極３０は、心臓内の心筋組織のマッピングの為に使用され得ることを察知すべきである。この使用で、内部電極は、例えば電位又は抵抗の形を取ることができる心臓内部の電気的な活動を検出する。検出された電気的な活動は、外科医によって分析される検出された活動を処理する外部制御器に送られる。
更に、多孔な拡張可能で−萎むことができる本体内部に配置された内部電極３０は、ペース取り信号を送る為に代わりとして、又は電気的な活動の検出と組み合わせて使用され得ることを察知すべきである。このように、内部電極３０は、ペースマッピング又は同調化マッピングを実行できる。
本発明の色々な特長は、次の請求の範囲に述べられている。Related applications
This application is a part of pending US patent application Ser. No. 07 / 951,728 filed on Sep. 25, 1992 under the name “Heart Mapping and Ablation System”. Partially pending US patent application Ser. No. 08 / 099,994 filed Jul. 30, 1993 under the title of “cardiac ablation catheter with large surface that flattened during introduction into the heart” It is.
Field of Invention
The present invention relates generally to electrode structures that are deployed within the body. More specifically, the present invention relates to an electrode structure that can be deployed in the heart to diagnose and treat heart disease conditions.
Background of the Invention
The treatment of cardiac arrhythmias requires electrodes that can create tissue damage with a variety of different geometric shapes and characteristics based on the specific physiology of the arrhythmia being treated.
For example, a conventional 8F diameter / 4 mm long cardiotomy electrode has a depth of about 0.5 cm and a width of about 10 mm and is 0.2 cm^ThreeRadio frequency energy can be transmitted to create damage in the myocardial tissue with a damaged volume that extends to These small and shallow injuries are needed in the sinus node to perform sinus node remodeling, or along the AV groove for various ancillary pathotomy or because of atrial spasm (AFL) Or along the slow zone of the tricuspid isthmus for AV nodal slow-passway incisions.
However, to remove the substrate of digestive organ tachycardia (VT), at a penetration depth deeper than 1.5 cm and a width greater than 2.0 cm, at least 1 cm.^ThreeIt appears that it requires significantly larger and deeper damage with a damaged volume of.
Furthermore, there remains a need to create damage with a relatively large surface area at shallow depths.
One solution that has been proposed to create a variety of damage features is to use different forms of ablation energy. However, the peripheral techniques surrounding microwave, laser, or ultrasonic ablation and chemical ablation have not been tested for this purpose.
It is known that employing a positive cooling method in connection with the transmission of DC or radio frequency ablation energy forces the interface between the electrode and tissue to be at a lower temperature. As a result, the hottest tissue temperature region is shifted into the tissue, which in turn shifts the tissue boundaries that are no longer viable by resection deeper into the tissue. A positively cooled electrode is used to transfer more ablation energy into the tissue compared to the same electrode that is not actively cooled. However, active cooling control is required to keep the maximum tissue temperature safely below about 100 ° C., which is known to cause tissue dehydration and tissue boiling.
Another solution that has been proposed to create greater damage in either surface area and / or depth is to use a substantially larger electrode than is commercially available. However, the larger electrodes themselves have some problems in size and operability and become heavy and prevent the safe and easy introduction of large electrodes into the heart through veins and arteries.
There is a need for a multi-purpose cardiotomy electrode that can selectively create lesions with different geometric shapes and characteristics. The multipurpose electrode will have the flexibility and operability necessary to allow it to be safely and easily introduced into the heart. Once deployed inside the heart, these electrodes, in a controlled manner, have enough energy to create the required treatment for large and deep injuries, small and shallow injuries, and large and shallow injuries. Has the ability to fire.
Summary of invention
The present invention is capable of causing ion transport of electrical energy without substantially spraying a liquid, and provides various porous electrode assemblies (assemblies) for use in a heated tissue ablation system.
A porous electrode assembly embodying features of the present invention has a wall whose outer side surrounds the inner region. A lumen carries the ion-containing medium into the inner region. One element couples the medium in the inner region to the electrical energy source.
According to one aspect of the present invention, at least a portion of the wall is composed of a porous material sized to prevent passage of macromolecules while allowing passage of ions contained in the medium. The wall thereby allows ion transport of electrical energy through the porous material to the outside of the wall.
According to another aspect of the present invention, at least a portion of the wall is a porous material sized to allow ions contained in the medium to pass therethrough without substantially liquid dispersion in the porous material. It is composed of Thereby, the wall allows ion transport of electrical energy through the porous material to the outside of the wall without substantially spraying liquid in the wall.
According to another aspect of the present invention, at least a portion of the wall is composed of a porous material dimensioned to allow ions contained in the medium to pass through the porous material to the outside of the wall. It enables ion transport of electrical energy. According to this aspect of the invention, the porous material has a bubble point value greater than the internal pressure.
According to another aspect of the invention, at least a portion of the wall is comprised of a hydrophilic porous material dimensioned to allow ions contained in the medium to pass through the porous material. Ion transport of electrical energy to the outside of According to this aspect of the present invention, the porous material has a bubble point value greater than the internal pressure, thereby allowing ion transport without substantial liquid spraying into the porous material. Get up.
Other features and advantages of the invention are set forth in the following description and drawings.
[Brief description of the drawings]
FIG. 1 is a plan view of a cardiac tissue ablation system having an expandable porous electrode structure embodying features of the present invention.
FIG. 2 is an enlarged side elevational view of a partially broken portion of the porous electrode structure used in connection with the system shown in FIG. 1, showing the electrode structure in its expanded configuration.
FIG. 3 is an enlarged side elevational view of the porous electrode structure shown in FIG. 2, showing the electrode structure in its deflated form.
FIG. 4 is a somewhat schematic side view, partially broken and further enlarged, of the porous electrode structure shown in FIG.
FIG. 5 is a porous electrode structure used in connection with the system shown in FIG. 1, showing the electrode structure in its expanded configuration due to the presence of an internal elongated flexure support structure. FIG. 5 is an enlarged side elevational view of the partial fracture of FIG.
6 is an enlarged side cross-sectional view of the porous electrode structure shown in FIG. 5, showing the electrode structure in its deflated form by manipulation of the outer sliding sheath.
FIG. 7 shows a porous electrode structure used in connection with the system shown in FIG. 1, showing the electrode structure in its expanded configuration due to the presence of an internal woven mesh support structure. FIG. 3 is an enlarged side elevational view of a partial body break.
FIG. 8 is a somewhat schematic enlarged view taken and expanded substantially along line 8-8 of FIG. 4, showing ion current density across the pores of the electrode body shown in FIG. is there.
FIG. 9 is a graph showing the relationship between the detected impedance and ion transport through the pores of the electrode body shown in FIG.
FIG. 10 shows a partial fracture of an alternative porous electrode structure used in connection with the system shown in FIG. 1, showing the electrode structure composed of porous foam in its expanded form. It is an enlarged side elevation view.
FIG. 11 is an enlarged side view of a porous electrode structure used in connection with the system shown in FIG. 1, wherein the pores of the structure are arranged in a bull's eye pattern at the end of the body. FIG.
FIG. 12 shows a porous electrode structure used in connection with the system shown in FIG. 1 in which the pores of the structure are arranged in compartments spaced circumferentially along the sides of the body. It is an enlarged side view.
FIG. 13 is a partially cutaway side view showing the use of multiple chambers for transporting liquid to the compartmental pore region shown in FIG.
FIG. 14 is an enlarged side elevational view of the porous electrode structure used in connection with the system shown in FIG. 1, further holding a non-porous electrode element.
FIG. 15 is an enlarged side cross-sectional view of a porous electrode structure that also holds electrode elements formed by wires that meander through the structure body.
FIG. 16 is an enlarged side elevational view of a porous electrode structure with internal pace distribution / detection electrodes.
Figures 17 and 18 graphically display tissue temperature profiles associated with porous electrode structures when operated at different conditions.
FIG. 19 is a diagram schematically showing a fixture and a mandrel for forming a hemispherical shape from a flat sheet of porous material to the end of the porous electrode body.
FIG. 20 is a side cross-sectional view of the fixture and mandrel shown in FIG. 19 in the process of forming the hemispherical end profile with a flat sheet of porous material.
FIG. 21 is an enlarged side cross-sectional view of the porous material sheet after forming the hemispherical end portion outer shape.
FIG. 22 is a schematic diagram of a finished fixture for forming the hemispherical shape of the proximal end of the porous electrode body from the pre-formed sheet shown in FIG.
FIG. 23 is an elevation view of the porous electrode body after it has been formed by the apparatus shown in FIGS.
FIG. 24 is used in place of the finished fixture shown in FIG. 22 to form the hemispherical shape of the proximal end of the porous electrode body from the pre-formed sheet shown in FIG. FIG. 3 is a diagrammatic representation of an expandable finishing fixture.
25 is an elevational view of the porous electrode body after being formed by the expandable fixture shown in FIG.
FIG. 26 is a schematic diagram of two pre-shaped hemispherical body portions of a porous electrode body before being joined together to a composite porous electrode body.
FIG. 27 is a side elevational view of a composite porous electrode body formed by joining the two hemispherical portions shown in FIG. 26 along a peripheral seam.
FIG. 28A is a side cross-sectional view showing the porous electrode body shown in FIG. 27 upside down to provide a peripheral seam inside the body to prevent direct contact with tissue.
28B is a side cross-sectional view of the porous electrode body shown in FIG. 27 after being turned over to provide a peripheral seam inside the body.
FIG. 29A is a side cross-sectional view of a porous electrode body formed after two hemispherical portions are joined along an axial seam and turned over to provide an axial seam inside the body.
FIG. 29B is a top view of the porous electrode body with the inverted axial seam shown in FIG. 29A.
FIG. 30 shows a porous structure formed by joining two hemispherical portions along a main axial seam, partitioning the body with an additional intermediate axial seam, and then turning the inner seam inside the body after turning over. It is a top view of an electrode main body.
FIG. 31A is an enlarged side cross-sectional view of a seam before two sides of a porous material are joined together to form an electrode body and the body that encloses the temperature sensing element in the sheet is turned over.
FIG. 31B is a side elevational view of the seam coupling body partially shown in FIG. 31A prior to turning over the body surrounding the temperature sensing element within the seam.
FIG. 31C is a side view of the side of the main body shown in FIG. 31B in which the seam and the signal wire of the temperature detection element are provided inside the main body after turning over.
FIG. 32A is a somewhat schematic representation of a porous electrode body that is formed from a regenerated cellulosic material by dipping using an expandable fixture.
FIG. 32B shows the dip-molded body formed in FIG. 32A after the expandable fixture is removed and the fixture with the steering assembly is attached to the end of the body and before it is turned over. .
FIG. 32C shows the dip-molded body with the end fixture and steering assembly shown in FIG. 32B after flipping.
FIG. 33 shows an exemplary porous body shaped into an elongated cylindrical outer shape with a radius varying along the length to form a distal region and a proximal neck region.
FIG. 34A shows another exemplary porous body shaped as an elongated cylindrical outer tube with a constant radius along its length.
FIG. 34B shows the tube shown in FIG. 34A with the end closed by a seam and a port tube for attachment to the catheter tube sealed to the proximal end.
The present invention may be embodied in several forms without departing from the spirit or essential characteristics. The scope of the invention is defined in the appended claims rather than the specific description set forth before the claims. Accordingly, all embodiments that fall within the spirit and equivalent scope of the claims are to be embraced by the claims.
DESCRIPTION OF PREFERRED EMBODIMENTS
FIG. 1 illustrates a tissue ablation system 10 that embodies the features of the present invention. The system 10 includes a flexible catheter tube 12 having a proximal end 14 and a distal end 16. The proximal end 14 includes a handle 18. The distal end 16 includes an electrode structure 20 that embodies the features of the present invention. The purpose of the electrode structure 20 is to transmit ablation energy.
As best shown in FIGS. 2 and 3, the electrode structure 20 has a body 22 that is expandable and can be deflated. The outer shape of the body 22 is changed between a deflated configuration (FIG. 3) and an expanded or expanded configuration (FIG. 2). In the preferred embodiment shown, liquid pressure is used to inflate and maintain the expandable body 22 in an expanded configuration.
In this configuration (see FIG. 2), the catheter tube 12 has an internal lumen 34 along its length. The distal end of the lumen 34 is open to the hollow interior of the body 22 which is expandable and can be deflated. The proximal end of the inner lumen 34 is in communication with the port 36 (see FIG. 1) of the handle 18. A liquid inflation medium (arrow 38 in FIG. 2) is carried through port 36 into lumen 34 under positive pressure. The liquid medium 38 fills the interior of the body 22 which is expandable and can be deflated. The liquid medium 38 is expandable by applying internal pressure-the body 22 that can be deflated from its deflated form.
With this feature, when the expandable-deflable body 22 is introduced into the venous structure, it takes a deflated low profile (ideally less than 8 French diameters, ie less than about 0.267 cm). Will be able to. Once placed in the desired location, the expandable-deflable body 22 is converted to a fully expanded form, eg, about 7-20 mm.
As shown in FIGS. 5-7, the structure 20 is normally open and deflated, if necessary, to increase the liquid medium pressure or to replace the internal force to maintain the body 22 in the expanded configuration. An internal support structure 54 can be provided. The shape of the internal support 54 can be changed. For example, from an assembly of flexible elongated elements 24 as shown in FIG. 5, an internal porous woven mesh or an open porous foam structure 26 as shown in FIG. Composed.
In these configurations (see FIG. 6), an expandable-deflable body 22 supported from the inside is added by an outer sheath 28 (see FIG. 6) that slides along the catheter tube 12. It is converted into a deflated form after the expansion medium is removed by an external compressive force. As shown in FIG. 6, forward movement of the sheath 28 advances it over the body 22 which can be expanded and deflated. The body 22 that is expandable and can be deflated is deflated within its sheath 28 to its lower profile. With the rearward movement of the sheath 28 (see FIG. 5 or 7), it is retracted away from the expandable-deflable body 22. When released from restraint by the sheath 48, the inner support structure 54 opens with spring force, returning the expandable-deflable body 22 to its expanded configuration and accepting the liquid medium.
As best shown in FIG. 4, the structure 20 further includes an internal electrode 30 made of a conductive material inside the main body 22. The material of the internal electrode 30 has both a relatively high conductivity and a relatively high thermal conductivity. Materials with these characteristics include gold, platinum and platinum / iridium. Precious metals are preferred.
An insulated signal wire 32 is connected to the electrode 30. A signal wire 32 extends from the electrode 30 through the catheter tube to an external connector 38 on the handle 18 (see FIG. 1). Connector 38 electrically connects electrode 30 to radio frequency generator 40.
In the preferred illustrated embodiment (see FIG. 1), they are combined with the radio frequency generator 40 either as a unit integrated into the controller 42 or as an independent alignment box. The controller 42 controls the supply of radio frequency energy to the electrode 30 according to pre-established criteria. Furthermore, details of this aspect of the system 10 will be performed later.
According to the present invention, the liquid medium 38 used to fill the body 22 is a conductive liquid. The liquid 38 forms a conductive path that carries radio frequency energy from the electrode 30. Relatedly, the main body 22 is made of a non-conductive thermoplastic material or elastic material having pores 44 at least on a part of its surface. The pores 44 of the porous body 22 (shown schematically in FIG. 4 in enlarged form for illustration purposes) allow the ablation energy from the electrode 30 to ion transport through the conductive medium 38 to the tissue outside the body. Let Preferably, the liquid 38 has a low resistance that reduces ohmic losses in the body 22 and thus the heating action of the ohms. In the preferred embodiment shown, the liquid 38 also at least partially serves as an expansion medium for the body.
The components of the conductive liquid 38 can be changed. In the preferred embodiment shown, the liquid 38 is composed of high osmotic brine having a saturated or near-saturated sodium chloride concentration of about 9% by weight per volume. High osmotic salt water has a resistance that is only about 5 Ω · cm lower than the blood resistance of about 150 Ω · cm and the myocardial tissue resistance of about 500 Ω · cm.
The component of the conductive liquid 38 is composed of a high osmotic pressure potassium chloride solution. This medium facilitates the desired ion transport, but the rate at which ion transport passes through the pores 44 needs to be monitored more closely to prevent potassium overload. It is desirable to keep the ion transport rate below about 10 mEq / min when a high osmotic potassium chloride solution is used.
As just described, the present system 10 is ideally suited for resection of myocardial tissue within the heart. In this embodiment, the surgeon moves the catheter tube 12 through the vena cava or artery into the ventricle while the expandable-deflable body 22 of the electrode structure 20 is in its low profile configuration. Yes. Once the expandable body 22 is expanded to its expanded configuration within the desired ventricle, the region with the pores 44 is expanded to contact the target region of intracardiac tissue. The
For the most part, depending on the mass concentration difference across the pore 44, the ions in the medium 38 will pass through the pore 44 due to diffusion caused by the concentration difference. Diffusion through the pores 44 continues as long as the concentration gradient is maintained across the body 22. Ions contained in the pores 44 provide a means for conducting current across the body 22.
Radio frequency energy is controlled by the controller 42 and delivered from the generator 40 to the electrode 30. When a radio frequency (RF) voltage is applied to the electrode 30, the current is carried by the ions in the pores 44. The ions move back and forth slightly while applying the RF frequency, but the RF current provided by the ions does not cause any net diffusion of ions, as occurs when a DC voltage is applied. This ionic motion (and current flow) in response to the applied RF field does not require liquid distribution of the media 38 through the pores 44.
The ions carry radio frequency energy through the pores 44 into the tissue facing the return electrode, which is typically an external patch electrode (forming a monopolar structure). Instead, the transmitted energy can pass through the tissue (forming a bipolar structure) to adjacent electrodes in the heart chamber. Radio frequency energy heats the tissue, mostly in electrical resistance, to form a therapeutic lesion.
The electrical resistance of the main body 22 has a great influence on the shape and controllability of the damage. It has been found that ablation with a device having a low resistance body 22 requires more RF power and causes deeper damage. On the other hand, a device with a high resistance body 22 provides more uniform heating and thus improved damage controllability. Since the body resistance is increased to generate additional heat, less RF power is required to reach the same tissue temperature after the same time interval. As a result, the damage generated in the high resistance body 22 is less deep than usual.
In general, a lower body 22 resistance value of less than about 500 Ω · cm will form a deeper damage shape. Similarly, a higher body 22 resistance of about 500 Ω · cm or more will form a shallower damage shape.
The electrical resistance of the main body 22 is controlled by specifying the pore size of the material, the porosity of the material, and the water absorption characteristics (hydrophilicity versus hydrophobicity) of the material.
Determination of pore size
The size of the pores 44 in the body 22 can be varied. Generally, pore sizes smaller than about 0.1 μm used for blood oxygen saturation, dialysis or ultrafiltration are used for ion transport according to the present invention. These small pores viewed by a high energy electron microscope constrain macromolecules, but allow ion transport through the pores in response to an RF field applied as described above. With smaller pore diameters, even if liquid spraying by pressure drive through the pores 44 is performed, ion transport is unlikely unless a relatively high pressure state is formed in the body 22.
The larger pore diameters commonly used for blood microfiltration are also used for ion transport according to the present invention. These larger pores seen by light microscopy techniques constrain blood cells but allow ions to pass through in response to the applied RF field. In general, pore sizes of less than 8 μm will prevent most blood cells from crossing the membrane. At larger pore diameters, pressure-driven liquid dispersal and accompanying transfer of macromolecules through the pores 44 are more likely to occur at normal filtration pressures on the body 22.
Larger pore sizes that can accommodate the blood cell elements formed are also used. However, the overall porosity and spreading rate and the occupation of blood cells within the pores of the body 22 must be taken into account as the pore size increases.
A conventional porous and biologically compatible membrane material used for blood filtration such as blood oxygen saturation, dialysis, and plasma export can serve as a porous body 22. Such a membrane material is made of, for example, regenerated cellulose, nylon, polycarbonate, polyvinylidene fluoride (PTFE), polyethersulfone, modified acrylic copolymer or cellulose acetate.
Instead, it is porous by weaving the material (such as nylon, polyester, polyethylene, polypropylene, fluorocarbons, small diameter stainless steel and other fibers) into a mesh with the desired pore size and porosity. Or the raw material which has a micropore is manufactured. It is advantageous to use a woven material because the woven material is very flexible because small diameter fibers are used to weave the mesh. Uniformity and consistency in pore size is also achieved by using a woven material.
The Spectrum Medical Industry Company (Houston, Texas) supplies the market with nylon and polyester woven materials with a pore size of 5 μm with a porosity of 2%. Stainless steel woven material with a porosity of 30% and a pore size of 3 μm is also available from Spectrum Medical Industries. Manufacturers like Tekoto also produce woven materials that meet the desired specifications. Woven materials with smaller pore dimensions are made on the basis of the material.
Woven mesh is manufactured by conventional techniques including square woven mesh or twill mesh. The square woven mesh is formed by the “vertical duplex” method. A twill mesh is formed by sending two fibers up and down. The material is woven into a three-dimensional structure such as a tube or sphere. Instead, the material is also woven into a flat two-dimensional sheet and formed into the desired three-dimensional form of the body 22 (by thermoforming, thermal bonding, mechanical deformation, ultrasonic welding, etc.).
The pore size is specified using a bubble point measurement method. The bubble point value is defined primarily as a function of the pore size (assuming the same water adsorption characteristics) and the pressure required to force the liquid through the membrane. . The standard for measuring the value of the bubble point is ASTM F316-80.
The pore size correlates with the expected liquid flow resistance of the membrane. As a general proposition, larger pores allow more liquid to flow through the pores at higher flow rates. Similarly, smaller pores limit the volume and rate of liquid distribution through the pores. At some point, the pores are small enough to effectively prevent liquid spreading except at very high pressures, while nevertheless ion transport can occur in the following situations: I am doing so.
It is preferable to reduce or substantially not spray the liquid through the pores 44. It may be advantageous for several reasons to restrict or substantially not distribute the liquid through the pores 44. First, it can limit salt and water overload conditions caused by transferring hyperosmotic solutions to the blood pool. This is especially true if the hyperosmotic solution contains potassium chloride as described above.
Furthermore, by restricting or substantially preventing liquid spraying through the pores 44, ion transport can be performed without confusion. By not being disturbed by the accompanying liquid distribution, a continuous virtual electrode 48 (see FIG. 8) is created at the intermediate surface between the body 22 and tissue in ion transport. The virtual electrode 48 efficiently transfers RF energy without the need for a conductive metal surface.
The psi bubble point value for a given porous material will help further shape the ion transport characteristics, thereby making the porous material compatible with tissue ablation.
If the value of the bubble point for a given porous material exceeds the pressure required to inflate the body 22 (ie, the expansion pressure), it does not promote pressure-driven fluid distribution through the material pores 44. In addition, the body 22 can be inflated with pressure into its expanded configuration. By specifying the material with a bubble point value that is greater than the pressure that causes the body to inflate, ion transport through the pores 44 is ensured without any additional liquid spraying through the pores 44. .
The value of the bubbling point with a much lower pressure to inflate the body would prevent the body 22 with porous material from reaching its targeted expansion configuration due to excessive liquid distribution through the pores 44. It will also show that.
On the other hand, the bubble point value of the porous material should not exceed the tensile strength of the porous material. By identifying this relationship between the bubble point value and the tensile strength, the application of the liquid will occur before the abnormally high pressure propagates, reducing the risk of the body 22 bursting.
Specifying the bubble point value mediates the use of larger pore size materials. Larger pore size materials present problems of swelling and excessive fluid distribution through the membrane.
Porosity identification
The arrangement of the pores 44 and the dimensions of the pores 44 determine the porosity of the main body 22. The porosity represents a space on the main body 22 that does not include material, is empty, or is composed of pores 44. The porosity expressed as a percentage represents the volume percentage of the main body 22 that is not occupied by the main body material.
For materials having a porosity greater than about 10%, the porosity P (in%) is determined as follows.
P = 100 (1-ρ_b/ Ρ_m)
In that case, ρ_bIs the density of the body 22 as determined by its weight and volume, and ρ_mIs the density of the material making up the body 22.
In order to obtain porosity for materials having a porosity of less than about 10%, a scanning electron microscope is used to obtain the number of pores and their average diameter. The porosity P (in%) is then obtained as follows:
P = Nπ (d²/ 4)
In that case:
N is the density of the pores and (P_n/ A).
P_nIs the number of pores in the body 22.
a is the total pore area (cm²).
π is a constant of 3.1416.
d is the average diameter (cm) of the pores.
The size of the porosity affects the liquid flow resistance of the main body 22 as described above. The equivalent electrical resistance of the main body 22 further depends on its porosity. A material with a low porosity has a high electrical resistance, whereas a material with a high porosity has a low electrical resistance. For example, when exposed to a 9% high osmotic solution (5 Ω · cm resistance), a material with a 3% porosity will have a blood or tissue electrical resistance (between 150 and 450 Ω · cm). Can have an electrical resistance comparable to
The distribution of pores 44 for a given porosity also affects the efficiency of ion transport. Given a porosity value, instead of fewer but larger pore arrays, many smaller pore array 44 is preferred. When there are a large number of small pores 44, the current density is distributed so that the current density in each pore 44 becomes smaller. Decreasing the current density causes ion flow of electrical energy to the tissue with minimal reduction due to resistive heat loss.
In addition, many smaller pore arrangements 44 are preferred in lieu of fewer but larger pore arrangements to help add a convenient large fluid flow resistance. Providing a large number of small pores 44 limits the rate at which liquid spray occurs through each pore 44.
The dynamic change in resistance across the body 22 is brought about by changing the diameter of the body 22 made of a porous elastic material such as silicon. In this configuration, when the elastic main body 22 creates a predetermined porosity in a relaxed state, the elastic main body 22 is made into a porous state by drilling pores of the same size into an elastic material. When the elastic body 22 is expanded, its porosity remains essentially constant, but the wall thickness of the body 22 decreases. Thus, as the diameter of the body 22 increases, the resistance across the body 22 decreases to reduce the wall thickness of the body 22 and increase the surface area. As the surface area of the body 22 increases by a factor of 2, the thickness of the body 22 decreases by a factor of 2, resulting in a resistance of 4
As a result, the desired damage form is identified according to the form of the body 22. Thus, by using the same porous body 22 to control the form of the body 22, small damage, shallow and wide damage, and wide and deep damage can be formed.
Preferably, the porous body 22 should have a consistent pore size and porosity over the desired ablation area. Without a consistent pore size and porosity, the difference in electrical resistance of the body 22 across the ablation region results in a higher temperature local region with a higher current density. If the difference in electrical resistance is large enough, the damage will not be useful for treatment because it will not extend to the desired depth or length. Furthermore, the low porosity non-uniform portion of the body 22 will suffer physical damage from itself as a result of localized heating effects.
Identification of water absorption characteristics
The porous material of the main body 22 is either hydrophobic or hydrophilic. However, the water absorption characteristics of the porous material also affect the electrical resistance of the material.
For materials of the same pore size and porosity, hydrophilic materials have a greater ability to transport ions of radio frequency energy without significantly increasing the liquid flow rate through the material. Ions floating in the medium are more likely to occupy the pores of the hydrophilic material even when there is no driving pressure exceeding the bubble point value of the material, compared to the hydrophobic material. The presence of these ions in the pores of the hydrophilic material has the ability to flow an ionic current without the need for fluid distribution through the pores. As a result, the pore size can be more easily reduced for hydrophilic materials, thereby increasing the bubble point value and minimizing liquid spreading without adversely affecting the current carrying capability by the desired ions. Furthermore, the relationship between porosity and resistance is more direct in the case of hydrophilic materials than in the case of hydrophobic materials.
Some forms of nylon (eg nylon 6 and nylon 6/6) are examples of hydrophilic materials with high water absorption suitable for use as porous electrodes. The example nylon identified in Example-3 below has a moisture absorption of 4.0% to 4.2% at 65% relative humidity and 20 ° C.
Nevertheless, conventional medical grade “balloon” materials such as PET and PeBax are hydrophobic. The ions in the medium are unlikely to occupy the pores of the hydrophobic membrane, and there is no driving pressure beyond the bubble point value of the material compared to the hydrophilic material. As a result, hydrophobic materials tend to require more liquid flow through the pores to carry ions into the pores, thereby allowing electrical energy to pass through the porous material. ing. For such materials, the expansion pressure of the body should exceed the bubble point value to allow effective ion transport.
In addition, due to the higher surface tension of hydrophobic materials that try to limit ion flow into the pores, hydrophobic materials also show a greater tendency to cause material breakage in the pores compared to hydrophilic materials. . A large potential difference across each pore in a hydrophobic material will cause separation of water molecules, dielectric breakdown of the membrane material, and localized overheating. Breakage is related to high temperature action, opening the pores depending on the material and burning the material around the pores, generally reducing the quality of the material. Furthermore, material breakage produces harmful tissue effects similar to CD excision, such as tissue charring.
Therefore, by changing the water absorption characteristic of the porous material from a more hydrophobic characteristic to a more hydrophilic sex characteristic, it is possible to offset undesirable electrical characteristics without changing the pore size and porosity. For example, incidents of material breakage due to high current density and potential drop in the pores are reduced by increasing the porosity of the material. However, the incident of material breakage can be reduced without changing the porosity by selecting a hydrophilic material such as regenerated cellulose or nylon 6 or nylon 6/6, which generally has high water absorption. Instead, more hydrophilic coatings and surface treatments are applied to materials with fewer hydrophilic lines. For example, some materials are immersed in a specially formulated hydrophilic coating and exposed to ultraviolet light to bond the coating to the material surface. The approach is particularly useful when conventional “balloon” material is used for the body 22 if the coating can withstand the ablation temperature without degrading.
Other measures are employed to offset other undesirable electrical features that are dependent on pore size, porosity and water absorption characteristics. For example, for larger pore materials, or when a porous hydrophilic material is used, the spreading rate is controlled by controlling the fluid pressure across the body 22.
Instead, for larger pore materials, or if a porous hydrophilic material is used, a material that increases the viscosity and thereby reduces the spreading rate is added to the hyperosmotic solution. Examples of materials added to increase the viscosity include ion-contrast (radiopaque) materials, non-ionic glycerol and concentrated mannitol solutions.
For example, the electrical performance of a woven material with larger pore sizes is aided by the addition of an ion radiopaque Kotrast material such as Renographin®-76. By adding a radiopaque material to the aqueous solution, the body 22 is viewed under fluorescent fluoroscopy (ultrasound cardiac fluoroscopy based on a Kotrust material). The flow resistance of the porous material is effectively increased by the increased viscosity of the medium.
The use of an ionic material that increases the viscosity eliminates the need to excessively increase the membrane resistance based on the concentration of the ionic material. The following Table 1 summarizes the results of in vitro experiments using ion radiopaque materials with a woven nylon 13.0 mm disk detector.

Table 1 shows that ion contrast media can achieve the same resistance results and reduce the power required to create the desired damage.
For either hydrophilic or hydrophobic porous materials, the system 10 can include a device that detects impedance in proximity to the intermediate surface of the body and tissue. As FIG. 9 shows, the impedance decreases with increasing liquid spray flow rate until reaching a limit point where the impedance is stable despite increasing spray rate. By detecting the impedance, the minimum flow rate R_MIN(Where the impedance is too high) and the maximum flow rate R_MAXThe spray flow rate can be controlled in between (there will be a salt or moisture overload condition beyond which it can occur).
The surface area of the electrode 30 immersed in a conductive medium within the body is increased to enhance ion transport. However, desirable features such as small shape wetting and overall flexibility of the body impose practical constraints on electrode dimensions.
When the electrode 30 is brought close to the pore 44 of the main body 22, the efficiency of ion transport through the conductive medium is also increased. In addition, structural features that exhibit a flexible, small, and deflated shape are practical limitations in considering this.
Formation of body 22
An expandable-deflable body 22 is formed around or around the glass mold. In this configuration, the external dimensions of the mold match the desired expanded configuration of the body 22 that is expandable and can be deflated. The mold is immersed in a predetermined procedure in the body material solution until the desired wall thickness is achieved. The mold is then etched away leaving a shaped expandable-deflable body 22.
Instead, the expandable-deflable body 22 is blow molded from a further extruded tube. In this configuration, the body 22 is sealed at one end using an adhesive or heat welding. The other end of the main body 22 remains open. An expandable and wilting body 22 is placed inside the mold. An expansion body, such as high pressure gas or fluid, is introduced from the open tube end. When the tubular body 22 is inflated to take a mold shape, it is exposed to heat. The shaped expandable-deflable body 22 is then withdrawn from the mold.
The porosity of the main body 22 is given either before or after molding by a carbon dioxide laser, excimer laser, YAG laser, high-power YAG laser, electron particle bombardment, or the like.
As previously described, coatings and surface treatments are also applied to make the surface more hydrophilic to improve the electrical characteristics of the body 22 for tissue ablation.
A porous material available on the market is also formed into the body 22. For such poorly bonded materials formed by chemical processes such as regenerated cellulose, the materials can be dipped (shown schematically in FIG. 32A and described later), injection molding, or It is chemically formed into a three-dimensional shape by changing the diameter and shape during extrusion.
For those materials that are bonded by thermal bonding, laser welding, ultrasonic welding and adhesives, these bonding and welding techniques are used to form a three-dimensional shape from the sheet in which the material is used. There are various ways to use. Fixtures and mandrels are used to mold the body 22 in connection with heat and pressure.
19 to 23 show a preferable method for forming the porous material 200 into a desired three-dimensional shape of the ablation main body 22. As shown in FIG. 19, the sheet 200 is placed in the forming cavity 202 on the fixture 204. The shape of the molding cavity 202 corresponds to the shape desired for the distal end of the body 22. In the illustrated embodiment, the shape is substantially hemispherical.
As FIG. 20 shows, the forming mandrel 206 pushes a portion 208 of the sheet 200 into the forming cavity 202. The shape of the molding mandrel 206 matches the hemispherical shape of the molding cavity 202. The mandrel 206 fits snugly into the cavity 202 with the material portion 208 sandwiched therebetween. Thus, a desired hemispherical shape is set on the material portion 208 by pressure. Either the molding mandrel 206 and / or the molding cavity 202 can be heated to provide an additional heat set to the blank portion 208 within the molding cavity 202. The pressure in the forming cavity 202 and optional heat shape the material portion 208 from a flat shape to the desired hemispherical shape (see FIG. 21).
The sheet with the previously molded portion 208 is removed from the fixture 204 and placed on the finished fixture 210 (see FIG. 22). The finish fixture 210 has a distal end 212 having a shape that matches the shape of the previously molded portion 208. Portion 208 fits over the distal fixture end 212.
Finishing fixture 210 is hemispherical in the illustrated embodiment, but includes a proximal end 214 having a desired shape at the proximal end of body 22. The sheet 200 is well creased around the proximal end 214 of the fixture 210.
The finish fixture 210 has a base region 216 around which the remaining blanks of the sheet 200 are gathered by overlapping pleats 218. The sheet 200 will thereby firmly conform to the overall shape of the fixture 210.
The finish fixture 210 is heated to help provide an additional heat set to the desired shaped sheet 200 of the body 22. A clamshell mold (not shown) is further secured around the fixture 210 to facilitate the molding process.
Now, the sheet material formed as the porous main body 22 (see FIG. 23) is slid off the fixture 210. The material pleats 218 collected around the base region 216 of the fixture 210 are bonded together, for example, by thermal bonding or ultrasonic welding. This forms a narrowed neck portion 220 in the body 22 to facilitate attachment of the body 22 to the distal end of the catheter tube.
Prior to pleating, the sheet edge 217 is cut into a portion to minimize the amount of material that collects during pleating. After pleating, excess material is folded back and bonded to the neck portion 220 and / or otherwise cut away to form a smooth transition between the neck portion 220 and the distal end 208. .
Instead, after removal from the fixture 204, the sheet 200 with the previously shaped portion 208 is spanned around the expandable fixture 222 (see FIG. 24). The proximal end 217 of the seat 200 is tightly fastened around the neck portion of the fixture 222 by a fastening member 223.
The fixture 222 is configured by a balloon (for example, made of Teflon material) that is expanded using a gas or a liquid into a shape required for the main body 22. Thereby, the sheet 200 is formed by a fixture 222 that expands to take a desired shape.
Before or during expansion of the fixture 222, heat is applied to the end 217 of the sheet 200 to soften the material and assist in the forming process. External pressure is also applied to the proximal end of the seat 200 to assist in creating the neck portion 220 with the desired reduced diameter. This further helps prevent the material from “bunching” at the proximal end 217.
The fixture 222 itself is heated by using a heated gas or liquid to further expand the fixture. The heat additionally heat sets the sheet 200 to the desired shape of the body 22. An outer clamshell mold (not shown) is also secured around the fixture 222 to facilitate the molding process. Instead, an outer shell made of a material such as glass that is etched away is used to give the desired final shape.
For any heating process used to mold the body 22 using either the fixture 204 or 222, a heat sink (not shown) is provided so that the pore size does not change significantly during the heating process. , Used to cool the previously molded end portion 208.
Instead, the effect of heating on the pore size is inferred and taken into account when forming the sheet 200 in the first example prior to forming into the body 22. For example, if the pores open during molding, the pores are made proportionally smaller during manufacturing to take into account the dimensional increase during molding. Thus, the desired pore size will eventually be achieved while the sheet is being formed into the body 22.
After molding, the fixture 222 is shrunk and pulled out (see FIG. 25). The molded body 222 remains.
In yet another alternative process (see FIGS. 26 and 27), the body 22 is formed by joining two pre-shaped portions 225 along a peripheral seam 224. In the illustrated embodiment, the portion 225 is shaped as a hemisphere as shown in FIGS. 19-21 by cutting away excess material around its periphery. The portion 225 is similarly preliminarily molded by a molding process based on the characteristics of the material.
The seam 224 that joins the two portions 225 is formed by thermal bonding, ultrasonic welding, laser welding, adhesive bonding, stitching, or the like based on the properties of the material. The bonding and stitching method employed is selected to ensure that the seam forms an airtight and liquid tight part. The tensile strength of the seam 224 should also exceed the bubble point value of the porous material.
Instead, two generally circular flat portions of porous material cut to size from the sheet are joined around their perimeter by seams without pre-forming. This creates a normally deflated disc that surrounds the open interior. During use, air introduces liquid into the open interior, thereby expanding the disk to the shape required for the body 22. The disk also surrounds an internal support structure 54 (schematically illustrated in FIGS. 5-7) that causes the disk to be in the desired shape.
Preferably, after joining the hemispherical or flat portions 225 with the seam 224, excess material that extends beyond the seam 224 is cut away. It should be noted that because the contact between the tissue and the somewhat roughened surface portion of the seam 224 causes trauma, the joined portion 225 is preferably turned inside out. Inversion, a seam 224 will be provided inside the body 22 (as shown in FIG. 28B), away from direct contact with the tissue.
As shown in FIG. 28A, the joined portion 225 opens a small hole 250 at one end 252, inserts a pull wire 254, attaches it to the other end 256, and pulls the other end 256 through the hole 250. Flipped by. Thus, the attached hemispherical portion 225 is turned inside out.
In the above embodiment, the peripheral seam 224 extends around the axis of the main body 22. Instead, the seam 226 extends along the axis of the body 22 (as FIGS. 29A and 29B show) and is any two or more portions that are pre-shaped into a flat or three-dimensional shape. 228 can be coupled to the body 22. A mating fixture (not shown), each supporting the body portion 228, is used to keep the portion 228 stationary while heat or ultrasonic energy is applied to create the seam 226.
As FIG. 30 shows, another axially extending seam 230 is also provided in the porous material sheet to partition the sheet, rather than joining the sheet to another sheet. . Further details of the partitioned porous electrode will be described later. For illustrative purposes, FIG. 30 shows a somewhat exaggerated view of the hemispherical protrusion of the compartment along the seams 226, 230.
Preferably, the resulting body 22 is turned over as just described so that an axially extending seam 226 or 230 is provided inside the body 22.
It should be appreciated that the portions 225 or 228 shown in FIGS. 26-30 need not necessarily be made from the same material, whether pre-shaped into a flat or three-dimensional shape. Materials with different porous features are joined together by seams as just described. Instead, the porous material is joined by a seam to a non-porous material that is either conductive or electrically insulating. As an alternative, the conductive material is bonded to the insulating material by a seam, and the electrode has a double side where one (conductive) contacts the tissue and the other (electrically insulating) is exposed to the blood reservoir. A body can be provided. Virtually any flexible material suitable for use in connection with the electrode body is bonded using the seam according to this aspect of the invention. It should further be understood that the number of portions joined together by the seam to form a composite electrode body can be varied.
Of course, various specific shapes are also selected. As shown in FIG. 2, the preferred shape is basically a spherical shape with a spherical contour at the end, and is symmetrical. However, asymmetric or non-spherical shapes are also used. For example, the expandable-deflable body 22 is formed with a flat end profile that gradually curves inward or creates a neck for attachment of the catheter tube 12. As will be explained later, an elongated cylindrical shape is also used as shown in FIGS. 33 and 34B.
FIG. 10 shows an alternative expandable-deflable porous body 50. In this embodiment, the body 50 is constructed from an open-pore foam that is typically shaped to take an expanded form. The electrode 30 is surrounded by the foam body 50. A high osmotic pressure liquid medium 38 is introduced into the foam body 50 while filling the open pores so that the desired ion transport of the ablation energy can be achieved as already described. Ion transport using the foam body 50 also provides an external porous skin 51 (shown on the right side of FIG. If so, it will be executed.
In this configuration, the sliding sheath is advanced along the catheter tube 12 (as previously described) to compress the foam body 50 into a deflated configuration. Similarly, when the sheath is retracted, the compressive force is removed. The foamed body 50, which has lost its sheath, springs open, returning the expandable-deflable body 50 to its expanded configuration.
In the preferred embodiment shown, the distal steering mechanism 52 (see FIG. 1) enhances the operation of the porous electrode structure 20 during and after deployment. The steering mechanism 52 can be changed. In the illustrated embodiment (see FIG. 1), the steering mechanism 52 has a rotating cam wheel 56 connected to an external steering lever 58 held by the handle 18. The rotating wheel 56 holds the proximal ends of the left and right steering wires 60. Wire 60 passes through catheter tube 12 with ablation energy signal wire 32 and connects to the left and right sides of an elastic flexible wire or leaf spring (not shown) adjacent to distal tube end 16. Further details and other types of steering mechanisms are shown in US Pat. No. 5,254,088 by Lanquist and Simpson, which is incorporated herein by reference.
As shown in FIG. 1, the leaf spring of the steering mechanism 52 is supported within the distal end 16 of the catheter tube 12. As shown in FIG. 1, when the control lever 58 is moved forward, one of the control wires 60 is pulled to deflect or curve the leaf spring, and the distal catheter end 16 and the electrode structure 20 are brought together. Bend or curve in the direction. When the control lever 58 is moved backward, the other control wire 60 is pulled to bend or curve the leaf spring 62 and, at the same time, bend or curve the distal catheter end 16 and the electrode structure 20 in the other direction. In addition, as FIG. 32C shows, the steerable leaf spring 268 is a portion of the end fitting 270 that is itself attached to the end of the porous body 22. In this configuration, the leaf spring 268 extends beyond the distal catheter end within the tube 272 inside the porous body 22. Steering wires 60, 62 attached to the leaf spring 268 also pass through the tube 272. The proximal end of leaf spring 268 is attached to a hub 274 attached to the distal catheter end 16.
In this configuration, the leaf spring 268 is curved in both directions in the main body 22 by the back and forth movement of the control lever 58 on the handle 18. The leaf spring 268 moves the end fitting 270 and deforms the porous body 22 in the direction in which the leaf spring 268 curves.
In either configuration, the steering mechanism 54 can be used regardless of whether the body that is expandable and can be deflated is in its deflated configuration or in its expanded configuration.
FIGS. 32A and 32B illustrate a preferred method of attaching the end fitting 270 and the leaf spring 268 to the porous body 22. In FIG. 32A, the porous body 22 is formed by immersing an expandable fixture 278 having a desired shape in a solution of regenerated cellulose 278. Details of such an expandable fixture 278 have already been described in another context and illustrated in FIGS. It should be appreciated that the porous body 22 can be formed in a variety of other ways, as already described.
As FIG. 32B shows, the fixture forms a submerged porous body 22 having a proximal neck portion 280 and a distal neck portion 282. After the body 22 is molded, the expandable fixture 276 is deflated and withdrawn, as also shown in FIG. 32B.
As FIG. 32B shows, the distal neck portion 282 is attached using, for example, an adhesive, or by adhesive bonding or thermal bonding, mechanical bonding, screwing, wrapping, or any combination thereof. A sleeve 288 is used to attach around the distal fitting 270.
The fixture 270 has a leaf spring 268 pre-attached thereto and associated components as previously described. When initially attached to the fixture 270, the proximal neck portion 280 of the body 22 is oriented in the direction opposite the leaf spring 268.
After attaching the distal neck portion 282 to the fixture 270, the body 22 is turned around the distal fixture 270 to cover the leaf spring 268 as shown in FIG. 32C. The proximal end of the leaf spring 268 is attached to a hub 274 supported by the distal catheter end 16. The inverted proximal neck portion 280 is then attached to the distal catheter end using a sleeve 286. The sleeve 286 is attached around the catheter tube in a variety of ways including adhesive bonding, thermal bonding, mechanical bonding, screwing, wrapping, or any combination thereof.
Various alternative methods of attaching the porous electrode body to the distal end of the catheter are a co-pending patent entitled "Trunk Element to Attach Tube and Wire to an Expandable-Wilting Structure" (Attorney Label 2456A-6). It is disclosed in the application.
As will be described in more detail later, the end fitting 270 can also serve as a non-porous conductive portion on the porous body 22. Similar fittings 270 are placed elsewhere on the porous body 22 for the same purpose.
A probe (not shown) is also attached to the end fitting 270 instead of or in combination with the leaf spring 268. From there, the probe extends inside the body 22 (following upside down) through the catheter tube 12 to a suitable push-pull controller (not shown) on the handle 18. The probe is movable along the axis of the catheter tube 12 and is pushed and pulled axially on the distal fitting 270, thereby extending or shortening the body 22.
Further details regarding the attachment of the end fitting to the electrode body can be found in the co-pending patent entitled “Expandable and Swivel Electrode Structure Steering or Manipulating at the End” (Attorney Label 2458A-4). Shown in the application.
33 and 33A / 33B are exemplary elongated cylindrical electrode bodies that are assembled to the distal fitting as shown and attached to the distal catheter end 16 as shown in FIGS. 32B and 32C. Is shown.
In FIG. 33, the body 290 is shaped into an elongated shape by extrusion, dipping or molding at different radii to form the distal and proximal neck portions 280 and 282. A suitable end fitting 270 (shown in phantom) is mounted within the distal neck portion 282 and the elongate body 22 is turned upside down as shown in FIGS. 32B and 32C. Complete the assembly. Proximal neck portion 280 is then attached to distal catheter end 16 as shown in FIG. 32C.
In FIGS. 34A and 34B, the body 22 is formed from a tube 292 of material formed by extrusion, molding or dipping with a uniform radius (as shown in FIG. 34B). In this configuration (see FIG. 34B), a seam 294 formed as previously disclosed closes the end of the tube 292. The proximal end of the tube 292 is sealed around the tubular bore 296 for attachment to the distal catheter end 16. Instead, the distal end of tube 292 is sealed around distal fitting 270 (shown in phantom in FIG. 34B). In the latter case, the tube 292 is turned over around the distal fitting 270 prior to attaching to the distal end 16 of the catheter.
The pattern of pores 44 defining the porous region of the body can be varied. Preferably, as schematically shown in FIGS. 2 and 3, at least one third of the surface portion of the expandable and wiltable body 22 is free of pores 44. Yes.
The absence of pores 44 on at least the proximal 1/3 surface portion of the expandable and wilting body 22 is desirable for several reasons. This part is usually not in contact with the tissue and as a result does not serve any purpose even though there is a virtual electrode boundary. Furthermore, this part also shows the smallest diameter. If electrically conductive, this portion will have an undesirably maximum current density. By leaving the proximal portion of the smallest diameter that is not normally in tissue contact free of pores 44, at or at the end portion of the expandable-deflable body 22 that will make tissue contact. Ensure maximum current density is distributed nearby.
If it is expected that the distal portion of the body 22 will be directed to end contact with the tissue and the ablation is expected to occur, the porous region is of course expandable—the tip of the distal end of the body 22 that can be deflated Should be directed around. For orientation facing this end, the porous region is composed of a continuous cap that covers the end portion 1/3 to 1/2 of the body 22, as shown in FIGS. However, if end contact with tissue is contemplated, in a preferred embodiment (see FIG. 11), the conductive porous region is in a concentric “bull's eye” pattern around the distal end of the body 22. Divided into separate energy transfer zones 62 arranged.
If it is expected that the side region of the body 22 will be reoriented to contact the tissue, the porous region will be circumferentially around the distal end 1/3 to 1/2 of the body. It is preferably partitioned into spaced axially elongated energy transfer zones 62 (see FIG. 12).
When the perforated region is composed of compartmentalized zones 62 on the body 22, a group of sealed air bubbles 64 (see FIG. 13) individually group the liquid 38 into each perforated region compartment 62. Used for transport. Each bubble 64 is individually in communication with a lumen 66 and receives the conductive liquid of one porous compartment 62 that it is responsible for. The multiple moods 64 also provide the ability to more specifically control the shape of the inflated body 22 by selectively inflating some, but not all, moods 64 with liquid. It is something that
Foam 64 may be molded separately and inserted into body 22, or they may be integrally formed while molding body 22 which is expandable and can be deflated.
As FIG. 12 shows, the compartmentalized porous zone 62 is also well adapted for use in connection with the wilting of the body 22 which can be further expanded and wilted. In this configuration, the non-porous electrode is comprised of a pleated or prayed region 68. In order to create these regions 68, the body 22 mold is such that the expandable and wilting material is formed slightly thinner, recessed or ribbed along the desired region 68. It has a pre-shaped surface shape. The expandable-deflable body 22 deflates around these pleated regions 68 while consistently and uniformly deflating the body 22 over itself. Thus, the resulting deflated shape is made more uniform and compact.
It should be appreciated that the retractable body 22 shown in FIG. 12 can also be used for other patterns of porous regions. The pleated region 68 is further provided with pores if necessary.
FIG. 14 shows an example of an expandable and deflated electrode structure 70 that performs a dual function. The structure 70 has an expandable and wilting main body 22 that houses the internal electrode 30 as previously described. The body 22 contains a conductive fluid 38 and also has one or more porous regions 62 that allow ion transport of electrical energy as just described.
The structure 70 shown in FIG. 14 also has one or more conductive regions 72 on the surface of the body 22. In one embodiment (as shown in FIG. 14), the holeless conductive region 72 is formed by sputtering, evaporation, ion beam coating, electroplating over the deposited seed layer, or processing thereof. It can be expanded by the combination of-consisting of a metal such as gold, platinum, platinum / iridium coated on the body 22 that can be deflated. Alternatively, the non-porous conductive area 72 is also composed of a thin foil fixed to the surface of the body. Further alternatively, the non-porous conductive region 72 is also comprised of a rigid attachment (such as the end fitting shown in FIG. 32C) supported by the porous body 22 at one or more locations. A signal wire (not shown) in the main body 22 is electrically connected to the non-porous electrode. The signal wire passes through the catheter tube 12 to connect to a connector 38 held by the handle 18.
In the preferred embodiment (see FIG. 15), the perforated conductive region 72 is comprised of an insulated signal wire 26 that is threaded through the body 22 at a desired point in the electrical connection through the interior of the body. ing. The electrical insulation at the end of the serpentine pierced wire 26 is removed, exposing the portion of the wire that is preferably planarized to serve as the conductive region 72. In the flattened region, it is fixedly attached to the main body 22 by the conductive adhesive 73. Adhesive 73 is also preferably applied to the area of body 22 through which wire 26 passes to seal it. The same signal wire 26 is snaked through the body 22 multiple times to form multiple regions 72 as needed.
For various methods of attaching the non-porous electrode 72 and associated signal wire to the electrode body 22 that can be expanded and deflated, see "Improved Electrical Connection of Electrode Structure" (Attorney Label 2458A-5). ) In a co-pending patent application.
Non-porous electrode 72 is used to detect electrical activity in myocardial tissue. The detected electrical activity is sent to an external controller that processes the potential for analysis by the surgeon. The process creates a map of potentials or negative events to identify potential arrhythmic lesions. Once identified with the non-porous electrode 72, the porous region 62 is used to deliver radio frequency energy as previously described to ablate the lesion.
Alternatively, or in combination with the detection of electrical activity, the non-porous electrode 72 is used to send a pace signal. Thus, the non-porous electrode can perform pace mapping or synchronized mapping.
Preferably (see FIG. 16), the surface of the body 22 carries an electrode 100 suitable for monopolar or bipolar detection or for pacing. Although these electrodes 100 are disposed on the inner surface of the body 22, their action for detection and / or pacing is not compromised because of the good conductive characteristics of the body 22.
Different electrode placements are used for monopolar or bipolar detection or for pacing. For example, a pair of 2-mm long and 1-mm wide electrodes are provided on the inner surface of the body 22. Connection wires 102 are attached to these electrodes 100. In order to prevent high osmotic pressure solutions from shorting these electrodes, it must be covered with an electrical insulator 104 (eg, epoxy or adhesive). Preferably, the distance between the electrodes is about 1 mm to ensure that a good bipolar action potential diagram is obtained. The preferred installation locations of these internal electrodes are the tip and center of the end of the structure 22. Furthermore, if multiple zones are used, it is desirable to provide electrodes between the ablation regions.
It is also desirable to provide an opaque marker 106 on the inner surface of the body 22 so that the surgeon can guide the device to the target site under fluoroscopy. Any high atomic weight material is suitable for this purpose. For example, platinum or platinum / iridium is used to make the marker 106. A preferable installation location of these markers 106 is the tip and center of the end portion of the structure 22.
Thereby, the expandable-defilable structure 70 shown in FIG. 14 combines the use of a “solid” non-porous electrode 72 with a “liquid” or porous electrode 62. An expandable-deflable structure enables the mapping of myocardial tissue for therapeutic purposes using one electrode action and the excision of myocardial tissue for therapeutic purposes using different electrode actions To do.
In an alternative embodiment, the non-porous electrode 72 of the structure 70 is used in series with the porous region 62 to deliver radio frequency energy to ablate tissue. In this configuration, the signal wires responsible for region 72 are electrically connected to generator 40 and carry radio frequency energy for transmission by one or more regions 72. At the same time, the internal electrode 30 receives radio frequency energy for transmission by the medium 38 through the porous body. By performing ion transport across the porous structure surrounding region 72, the effective surface area of the ablation electrode is expanded.
In this example, the expandable-defilable structure 70 shown in FIG. 14 combines the use of an electrode 72 having a first effective surface area for detection and mapping. Yes. The first effective surface area is selectively increased for ablation purposes where ablation is performed by ion transport of hyperosmotic liquid across the porous structure surrounding electrode 72.
If liquid dispensing is performed through the pores, the internal electrodes 30 are not required to increase the effective electrode surface area of those regions. The liquid dispersion of ionic media through the pores, which occurs when the regions transmit radio frequency energy, itself serves to increase the effective transmission surface area of the regions 72. However, if ion transport is performed substantially without liquid spraying, the radio frequency energy may be transmitted using the internal electrode 30 simultaneously with the radio frequency being supplied to the external region 72 for transmission. It seems to be advantageous.
It should also be noted that in this embodiment, regions 72 are themselves made from porous conductive material. In this way, ion transport can take place across the region 72 itself.
As previously described (see FIG. 1), the controller 32 preferably controls the delivery of radio frequency ablation energy from the generator 30 to the electrodes supported within the body 22. In the preferred embodiment (see FIG. 2), the porous electrode structure 20 holds one or more temperature sensing elements 104 connected to the controller 32.
The temperature detection element 104 is in the form of a thermistor, a thermocouple, or the like. The temperature detection element 104 is in thermal conductive contact with the exterior of the electrode structure 20 and detects the condition of the tissue outside the structure 20 during ablation.
The temperature detected by the temperature detection element 104 is processed by the controller 32. Based on the input temperature, the controller 32 adjusts the time and power level of radio frequency energy transfer by the electrode 30 to achieve the desired damage pattern and other ablation purposes.
A co-pending patent entitled "Enhanced Electrical Connection for Electrode Structure" (Attorney Label 2458A-5) for various ways of incorporating temperature sensing element 104 into an expandable-deflable electrode body Explained in the application.
As shown in FIGS. 31A, 31B and 31C, the temperature sensing element 104 is further disposed near or within a seam 258 that joins the sheets 260, 262 together to the body 22. The formation of such a seam 258 has already been described and is further illustrated in FIGS.
As shown in FIGS. 31A and 31B, each temperature sensing element 104 is placed on one sheet 260 and then covered by the other sheet 262. The two sheets 260, 262 are then joined together to form the body 22. A seam 258 surrounds the detection element. The signal wire 264 of each detection element 104 extends where there is no seam 258 to the outside of the seats 260, 262, as shown in FIGS. 31A and 31B.
As previously described, the body 22 is preferably turned over (see FIG. 28A). As shown in FIG. 31C, both the seam 258 and the enclosed signal wire 264 are provided inside the main body 22 by turning over. Signal wire 264 is threaded through neck portion 266 for connection to controller 32.
Instead of or in addition to the temperature sensing element 104, the pace electrode and the sensing electrode can also be enclosed within a seam 258 of the inverted electrode body 22, as shown in FIGS. 31A-31C. . In such a configuration, after flipping, it is preferred to position the electrodes so that they are placed in a seam close to the material surface where the intended contact between the body material and the tissue takes place.
Further details on the use of many ablation energy transmitters controlled using a number of temperature sensing elements were filed on August 8, 1994, and “ Co-pending US patent application Ser. No. 08 / 286,930 entitled “Systems and Methods for Control”.
Example-1
In vitro analysis
Three electrode configurations were analyzed:
(1) Consists of dialysis tubing made from regenerated cellulose (manufactured by Spectra) with a molecular weight cut-off of 12,000-14,000 daltons and using 9% saline solution as the internal liquid medium A porous, expandable and wilting electrode structure according to the invention having a 13 mm disc-shaped body;
(2) a sputtered platinum disc-shaped electrode body having a diameter of 13 mm;
(3) An expandable-deflable hemispherical electrode structure made of aluminum foil having a diameter of 10 mm.
In the region where the high temperature condition existed, the thermistor was embedded by 0.5 mm in the tissue of the animal (sheep). With respect to the electrodes (1) and (2), the maximum temperature was generated at the edge of the disk-shaped body due to the edge heating effect. For electrode (3), the highest temperature occurred at the tip of the end of the hemispherical body with the highest current density.
The radio frequency electromagnetic power was adjusted to maintain the thermistor temperature at 60 ° C, 70 ° C, 80 ° C or 90 ° C. All lesion diameters were measured based on a 60 ° C. isotherm that marks tissue decolorization.
The following Table-2 lists the results observed in the in vitro analysis.

The porous electrode structure (1) was minimally affected by convective cooling compared to a regular metal ablation electrode such as electrode (2). This has been found by observing that there is no need to change the power to change the fluid flow rate during a particular damage occurrence and to maintain the thermistor temperature for the porous electrode structure.
Table-2 shows that the porous electrode structure created at least as much damage as the metallized electrode structure when adjusting power based on tissue temperature. The porous electrode structure had a reasonable impedance level as compared with the metal-coated electrode structure.
This example shows that the porous electrode structure can create damage deeper than 1.0 cm in a controlled manner to ablate epicardial, endocardial or intramural substrates.
The dialysis tube forming the porous electrode structure has high water absorption characteristics. Dialysis tubes become significantly more flexible when exposed to water. The molecular weight cut-off was 12,000 to 14,000 daltons. Larger or smaller molecular weight cut-offs are available from Spectrum. The conversion from the molecular weight cutoff to the predicted pore size for the dialysis tubing being tested is 0,000,000,000, as obtained from the medical instrument booklet 94/95 (p10) entitled “Dialysis / Ultrafiltration”. Equal to .01 μm, 50,000 Dalton equals 0.004 μm, 10,000 Dalton equals 0.0025 μm, and 5,000 Dalton equals 0.0015 μm.
Dialysis tubes have hydrophilic characteristics and high porosity despite their small pore size. As a result, the bubble point value is very high and the resistance is low enough that no liquid flow is required during the delivery of radio frequency energy.
Example-2
Finite element analysis
A three-dimensional finite element model was created for a porous electrode structure having an elongated shape with a total length of 28.4 mm, a diameter of 6.4 mm and a body wall thickness of 0.1 mm. A 0.2 mm diameter metal wire was extended inward along the length of the body to serve as an internal electrode connected to the RF energy source. The body was filled with a 9% hyperosmotic solution having an electrical resistance of 5.0 Ω · cm. The porous body of the structure was made as an electric wire. Close contact with the heart tissue was ensured along the entire length of the electrode body lying in the plane below the electrode. Contact with blood was ensured over the entire length of the electrode body lying in the upper plane of the electrode. The blood and tissue regions had resistances of 150 and 500 Ω · cm, respectively.
Analysis was performed on the electrode body based on resistances of 1.2 k-Ω · cm and 12 k-Ω · cm.
Table 3 shows the maximum tissue temperature depth when RF ablation power is applied to the porous electrode at various power levels and at various resistance levels relative to the porous body of the electrode.

FIG. 17 shows the temperature profile when the electrode is energized at 58 watts for 240 seconds when the porous body has a resistance of 1.2 k-Ω · cm. The depth of the 50 ° C. isotherm in FIG. 17 is 1.4 cm.
FIG. 18 shows the temperature profile when power is applied to the electrode at 40 watts for 240 seconds when the porous body has a resistance of 12 k-Ω · cm. The depth of the 50 ° C. isotherm in FIG. 18 is 1.0 cm.
In all cases, the maximum temperature is located at the intermediate surface between the tissue and the edges of the elongated porous structure. This is the preferred location of the temperature sensing element for the elongated shape of the porous body is at each end edge of the body. Preferably, each edge should hold at least one temperature sensing element, and the multiple temperature sensing elements are arranged on diagonally opposite sides to ensure that at least one of them faces the tissue. Should be.
The data further shows that the hottest regions, as observed with metal surface electrodes, are not moved deeply into the tissue. The hottest region is consistently present on the intermediate surface of the tissue-electrode body for direct detection. Although somewhat larger differences are encountered in certain other aspects of instrumentation, this feature is theoretically zero if the difference between the detected temperature and the virtually hottest tissue temperature is possible. Reduce to ℃.
A porous electrode body with a higher resistance body (see FIG. 18) is more uniform when viewed compared to a porous body with a lower resistance value (see FIG. 17). Temperature profile was generated. Less heat was required to reach the same maximum temperature due to the additional heating heat generated at the tissue-electrode body interface by increasing the resistance of the electrode body. The result was a reduced damage depth.
As previously described, by selecting the resistance of the body 22, the surgeon can greatly affect the lesion shape. Using a low resistance body 22 results in deeper damage and vice versa. Table 4 below illustrates the relationship between body resistance and damage depth based on empirical data.

Due to the reduced thermal conductivity of the porous electrode structure, damage formation is expected to be less sensitive to dynamic blood flow conditions around the electrode as compared to non-porous metal surface electrodes. The delivery of ablation energy through the porous electrode is more precisely controlled to obtain the desired damage characteristics, especially when shallow atrial damage is required.
The following Table-5 is based on empirical data and demonstrates a decrease in sensitivity of the porous electrode structure to convective cooling conditions due to changes in blood flow rate.

Using a porous electrode structure provides structural advantages. That is, placing the metal conductive shell on the outside of the expandable body that can be deflated isolates the adhesion problems that can occur as a related event. The porous electrode structure also avoids potential problems that can cause tissue adhesion to external conductive materials.
In addition to these structural advantages, the temperature control of the ablation operation is improved. When using conventional metal electrodes to ablate tissue, the tissue-electrode interface is cooled convectively by the surrounding blood flow. Due to these convective cooling effects, the region of maximum tissue temperature is located deeper in the tissue. As a result, the temperature condition detected by the sensing element associated with the metal electrode element does not directly reflect the actual maximum tissue temperature. Under this circumstance, the maximum tissue temperature condition must be inferred and predicted from the actual detected temperature. The porous electrode structure 20 or 70 is used to minimize the convective cooling of the tissue-electrode interface by the surrounding blood flow. As a result, the highest temperature region will be located at the intermediate surface between the tissue and the porous electrode. As a result, the temperature condition detected by the sensing element associated with the porous electrode element will more accurately reflect the actual highest tissue temperature.
Example-3
In vitro experiments were performed to compare hydrophilic materials (Hphi) versus hydrophobic materials (Hphb) with items using them as porous tissue excision elements. Table 6 summarizes the results.

Table 6 shows that the pore size is reduced using a hydrophilic material, thereby minimizing or stopping liquid spreading through the porous material while still allowing ion transport through the membrane. Prove that.
The hydrophobic porous material makes it possible to realize a high-resistance porous electrode. On the other hand, the hydrophilic porous material makes it possible to realize a low-resistance porous electrode.
Get the desired damage characteristics
As the above table proves, the same expandable and wilting porous electrode structure 20 can selectively form either wide and shallow or large and deep damage. Various methodologies are used to control the supply of radio frequency energy to achieve this result.
A. D50C action
In one exemplary embodiment, the controller 2 is located below the tissue-electrode interface from the surgeon to (i) the boundary depth between the tissue where the desired damage is viable and non-viable. An input 300 (see FIG. 1) that receives the desired treatment result in the area of widening and / or (ii) the highest tissue temperature that occurs within the lesion between the tissue-electrode interface and the boundary depth. The controller 42 has.
The controller 42 further includes a processing element 302 (see FIG. 1) that maintains a function that correlates the observed relationship between the damage boundary depth, the ablation power level, the ablation time, the actual subsurface tissue temperature, and the electrode temperature. )have. The processing element 302 compares the desired treatment result with the function and selects operating conditions based on the comparison to achieve the desired treatment without exceeding the actual or predicted subsurface tissue temperature described above. Achieve results.
The operating conditions selected by the processing element 302 include controlling the ablation power level, limiting the ablation time to the selected target ablation time, limiting the ablation power level to receive the maximum ablation power level, and / or region 44 and tissue. Various aspects of the ablation process can be controlled, such as the orientation of the porous region 44 of the body 22, including the prediction of the desired percentage contact between the body and the body. The processing element 302 relies on a temperature sensor held by or otherwise associated with the structure 20 that can expand and deflate into the tissue to detect the actual maximum tissue temperature. . Instead, the processing element 302 can predict the maximum tissue temperature based on operating conditions.
In a preferred exemplary embodiment, the electrode structure 20 holds at least one temperature detection element 10 that detects the instantaneous local temperature (T1) of the hot mass in the region 44. The temperature T1 at a certain time is a function of the power supplied to the electrode 30 by the generator 40.
The characteristics of the damage are expressed by a representation of the depth below the tissue surface in an isothermal region of 50 ° C. called D50C. Depth D50C is the physical characteristics of the porous region 44 (ie, its electrical and thermal conductivity, resistance, and dimensions); the percentage contact between the tissue and the porous region 44; and the local temperature of the hot mass in the region 44 It is a function of T1; the magnitude of RF power delivered by the internal electrode 30 (P); and the time (t) that the tissue is exposed to RF power.
For the desired damage depth D50C, further consideration of safety will constrain the selection of optimal operating conditions among the operating conditions listed in the matrix. The principal safety constraints are the highest tissue temperature TMAX and the highest power level PMAX.
The maximum tissue temperature condition TMAX is high enough to cause deep and wide damage (generally between about 85 ° C and 95 ° C) but safely below about 100 ° C and can cause tissue dryness and tissue microrupture. It is within the known temperature range. It has been recognized that TMAX occurs at some distance below the electrode-tissue interface between the interface and D50C.
The highest power level PMAX takes into account the physical characteristics of the internal electrode 30 and the power generation capability of the RF generator 40.
These relationships are observed empirically and / or by computer models under controlled actual and simulated conditions, as the above examples illustrate. The D50C action for a given porous region 44 can be expressed in all or some of the expressions listed in the matrix of values above and their relational expressions derived from empirical data and / or computer modeling. The
Processing element 302 has the operating conditions of this matrix in memory to define the D50C temperature boundary function for t = 120 seconds and TMAX = 95 ° C. and for a series of other operating conditions as described above. .
The surgeon further uses the identification code described above to identify features of the structure 20, and to set the desired maximum RF power level PMAX, the desired time t, and the desired maximum tissue temperature TMAX. 300 is also used.
Based on these inputs, processing element 302 compares the desired treatment result with a function defined in the matrix. The generator 42 controls the function variables to select the operating conditions so as to achieve a desired treatment result without exceeding the TMAX.
Thus, this configuration allows the surgeon to “dial and damage” formation by effectively identifying the desired D50C.
Furthermore, the details of eliciting the D50C action and its use to obtain the desired damage pattern are described in “Systems and Methods for Obtaining Desired Damage Characteristics During Body Tissue Resection” incorporated herein by reference. By name, it can be found in pending US application Ser. No. 08 / 431,790 filed May 1, 1995.
B. Comparted area: Duty cycle control
Various RF energy control methods are shown in FIG. 11 (the bull's eye pattern spaced apart in the axial direction of the zone) and in the partitioned porous pattern shown in FIG. 12 (the zones spaced apart in the circumferential direction). Used in conjunction with. For illustrative purposes, 44 (also referred to as the electrode area) is labeled E (J), but J represents a zone 44 (J = 1 to N).
As previously described, each electrode region E (J) has S (J) where J represents a certain zone 44 and K represents the number of temperature sensing elements (K = 1 to M) on each zone. , K) at least one temperature detection element 104.
In this mode, the condition is set via an appropriate power switching interface to supply RF power with multiple pulses of duty cycle 1 / N.
The amount of power (P) delivered to each individual electrode region E (J) with pulsed power supply_{E (J)}) Is expressed as follows:
P_{E (J)}αAMP_{E (J)} ²× DUTYCYCLE_{E (J)}
In that case:
AMP_{E (J)}Is the amplitude of the RF voltage sent to the electrode region E (J),
DUTYCYCLE_{E (J)}Is the duty cycle of the pulse expressed as:
DUTYCYCLE_{E (J)}= TON_{E (J)}/ (TON_{E (J)}+ TOFF_{E (J)})
In that case:
TON_{E (J)}Is the time during which the electrode region E (J) emits energy during each pulse period;
TOFF_{E (J)}Is the time during which the electrode region E (J) does not emit energy during each pulse period.
Formula (TON_{E (J)}+ TOFF_{E (J)}) Represents a pulse period for each electrode region E (J).
In this mode, the generator 40 is 1 / N duty cycle (DUTYCYCLE) for each electrode area._{E (J)}) Can be generated intensively
The generator 40 in turn generates successive power pulses for adjacent electrode regions so that the end of the previous pulse duty cycle slightly overlaps the beginning of the next pulse duty cycle. This overlap in the duty cycle of the pulse ensures that the generator 40 can continuously supply power without being caused by the release circuit during pulse switching between electrode regions of continuous interruption. is there.
In this mode, the temperature controller 42 controls each electrode region (AMP_{E (J)}) Of the ablation energy, which is individually adjusted to the amplitude of the RF voltage and is controlled by the generator 40 and delivered to each electrode region during the duty cycle_{E (J)}Switch individually.
In this mode, the generator 40 cycles through successive data acquisition sampling periods. During each sampling period, the generator 40 generates an individual sensor S (J, K) and a temperature code TEMP (J) (S (J, K) detected by the detection element 104 as output by the controller 42. K), which is the highest).
If more than one sensing element 104 is associated with a certain electrode area (eg, when a sensing element located at an edge is used), the controller 42 will detect all detected temperatures for a certain electrode area. And the highest detected temperature constituting TEMP (J) is selected from these.
In this mode, the generator 40 compares the temperature TEMP (J) locally detected at each electrode E (J) during each data acquisition period to the set point temperature TEMPSET established by the surgeon. On the basis of this comparison, the generator 40 determines the DUTYCYCLE for that electrode area and all other electrode areas._{E (J)}The amplitude AMP of the RF voltage supplied to the electrode region E (J) while maintaining_{E (J)}Change the set point temperature TEMP_SETGenerate and maintain TEMP (J).
The set point temperature TEMPSET can vary depending on the surgeon's judgment and empirical data. A typical set point temperature TEMP for cardiac resection is typically in the range of 40 ° C. to 95 ° C., although 70 ° C. is a typical preferred value.
Generator 40 is AMP_{E (J)}As a method for controlling, a proportional control method, a proportional integral derivation (PID) control method, or an ambiguous theoretical control method can be adopted.
For example, using a proportional control method, if the temperature detected by the first sensing element is TEMP (1)> TEMP_SETIf so, the amplitude AMP of the RF voltage to which the control signal generated by the generator 30 is applied to the first electrode region E (1)_{E (1)}On the other hand, the duty cycle DUTYCYCLE for the first electrode region E (1)_{E (1)}Keep the same. If the temperature detected by the second detection element TEMP (2) is TEMP (2) <TEMP_SETIf so, the control signal of the generator 30 is the amplitude AMP of the pulse applied to the second electrode region E (2)._{E (2)}On the other hand, the duty cycle DUTYCYCLE for the second electrode region E (2)_{E (2)}Duty cycle DUTYCYCLE_{E (1)}And keep it the same. If the temperature detected by a detection element is the setpoint temperature TEMP_SETIf so, no change in the amplitude of the RF voltage is caused for the associated electrode region.
The generator 40 continuously processes the voltage difference input during successive data acquisition periods to generate an AMP in each electrode region E (J)._{E (J)}, While maintaining the collective duty cycle the same for all electrode regions E (J). Thus, the mode maintains the desired uniform state of temperature over the length of the ablation element.
Using proportional-integral-derivative (PID) control techniques, the generator can change not only the instantaneous changes that occur in a sampling period, but also the changes that occurred in the previous sampling period and these changes over time. The ratio is also taken into account. Thus, using PID control techniques, a greater percentage of the difference has changed since the previous sampling period, based on whether the difference is larger or smaller compared to the previous instantaneous difference. Based on whether it is made smaller or smaller, the generator is TEMP (J) and TEMP_SETReacts differently to some proportionally large instantaneous difference between
In addition, for details on controlling the amplitude / collective duty cycle individually for partitioned electrode regions based on temperature detection, see “Multiple Temperature Detection Elements” incorporated herein by reference. System and Method for Controlling Tissue Ablation ", which can be found in pending US application Ser. No. 08 / 439,824, filed May 12, 1995.
C. Comparted area: Disabling at differential temperature
In this control mode, the controller 42 is the highest (TEMP for that phase) at the end of each data acquisition phase._SMAX) To select the detected temperature. The controller 42 also has a minimum (TEMP_SMIN) Is selected for that phase.
The generator will select the highest detection temperature TEMP selected._SMAXSelected high setpoint temperature TEMP_HISETCompare with In comparison, a proportional, PID or fuzzy theoretical control method is used to generate a control signal that centrally adjusts the amplitude of the RF voltage for all electrode regions.
In proportional control implementation plan:
(I) TEMP_SMAX> TEMP_HISETIf so, the control signal intensively reduces the amplitude of the RF voltage sent to all regions.
(Ii) TEMP_SMAX<TEMP_HISETIf so, the control signal intensively increases the amplitude of the RF voltage sent to all regions.
(Iii) TEMP_SMAX= TEMP_HISETIf so, there is no change in the amplitude of the RF voltage sent to all regions.
The generator detects the temperature TEMP for the purpose of amplitude control._SMAX, TEMP_SMINIt should be appreciated that any one of or between the temperatures can be selected to compare this temperature condition with a preselected temperature condition.
The generator has a certain local temperature TEMP (J) and TEMP_SMINThe power supply to the area is controlled based on the difference between them. In this execution, the locally detected temperatures TEMP (J) and TEMP_SMINThe difference between is calculated and this difference is compared with the selected set point temperature difference ΔTEMPSET. Therefore, a control signal for controlling power supply to the electrode region is generated.
If the local temperature TEMP (J) for a certain electrode region E (J) is ΔTEMP_SETSame as (or if TEMP (J) -TEMP_SMIN≧ ΔTEMP_SETΔTEMP)_SETLarger and lowest detected temperature TEMP_SMINIf it exceeds, the generator cuts off the power supply to a certain predetermined area E (J). TEMP (J) -TEMP_SMIN<ΔTEMP_SETIn this case, the generator returns the power supply to the certain predetermined area E (J).
Instead, TEMP (J) and TEMP_SMINInstead of comparing the TEMP_SMAXAnd TEMP_SMINCan be compared. TEMP_SMAXAnd TEMP_SMINThe difference between and a predetermined amount ΔTEMP_SETIs greater than or equal to_SMINBlock all areas except the one where exists. TEMP_SMAXAnd TEMP_SMINThe temperature difference between_SETIf so, the generator 30 restores these areas.
Further details on using differential temperature disabling can be found in the name of “Systems and Methods for Controlling Tissue Ablation Using Multiple Temperature Sensing Elements”, incorporated herein by reference. See pending US patent application Ser. No. 08 / 286,930, filed 8 months ago.
D. PID control
For the porous electrode structure, the actual detected temperature condition can be used as the maximum tissue temperature TMAX instead of the expected temperature with minimal action of convective cooling by the blood pool. Thereby, such structures also use their own proportional integral derivation (PID) control technique. An illustrative PID control technique available in connection with these electrode structures is named “Tissue Heating and Ablation System and Method Using Time-Setpoint Temperature Curves for Monitoring and Control”, June 1994. This is disclosed in pending US patent application Ser. No. 08 / 266,023 filed on Jan. 27.
Finally, it should be appreciated that the internal electrode 30 disposed within the porous expandable-deflable body can be used for mapping myocardial tissue within the heart. With this use, the internal electrodes detect electrical activity inside the heart, which can take the form of, for example, a potential or resistance. The detected electrical activity is sent to an external controller that processes the detected activity that is analyzed by the surgeon.
In addition, it has been found that the internal electrode 30 disposed within the porous expandable-deflable body can be used as an alternative or in combination with electrical activity detection to send pace signals. Should. In this way, the internal electrode 30 can perform pace mapping or synchronized mapping.
Various features of the invention are set forth in the following claims.

Claims (14)

An expandable and contractible body having an inner surface and an outer surface, surrounding the inner region and having a wall capable of holding an ion-containing medium under pressure, and disposed within the inner region, the medium and an electrical energy source An electrode assembly for a catheter comprising electrode elements arranged so as to be connectable,
For a catheter in which at least a part of the wall is composed of a porous part having a bubble point larger than the expansion pressure of the main body, and ion transport of electric energy from the electrode element to the outer surface of the wall can be performed via the medium. Electrode assembly. The electrode assembly according to claim 1, wherein the wall includes a terminal region and a proximal region, and the porous portion occupies more of the terminal region than the proximal region. The assembly of claim 2, wherein there are no holes in at least one third of the proximal region. 2. The electrode assembly according to claim 1, wherein the porous portion includes at least a first porous region and a second porous region which are arranged separately via a third region where no pore exists. 5. An electrode assembly according to claim 4, wherein the first and second porous regions are arranged circumferentially separated through the third region about an axis defined by a catheter. The electrode assembly according to claim 1, further comprising at least one temperature sensing element held on the wall and in thermal contact with the outer surface. 2. The electrode assembly according to claim 1, wherein at least a part of the wall is made of a conductive material. The electrode assembly according to claim 7, wherein the porous portion of the wall is adjacent to the conductive material. The electrode assembly according to claim 7, wherein the conductive material is porous, and ions contained in the medium can pass through the conductive material. The electrode assembly according to any one of claims 1 to 9, wherein the porous material is hydrophilic. The wall
A first body portion formed from a first flexible material;
A second body portion formed from a second flexible material;
2. The electrode assembly according to claim 1, further comprising a seam that joins the first and second main body portions to form a composite. The electrode assembly according to claim 11, wherein the first flexible material has higher conductivity than the second flexible material. The electrode assembly of claim 12, wherein the second flexible material is substantially non-conductive. The electrode assembly according to claim 1, further comprising an element that is held on the inner surface of the wall and that is disposed so as to be in contact with the outer surface of the wall and capable of measuring the electrical activity of a living tissue.

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