JP3682270B2 - Optical waveguide type glucose sensor - Google Patents

Description

メッシュ状導電性薄膜１０は、例えばチタン薄膜等から作られる。
【００１８】
次に、前述した光導波路型グルコースセンサの製造方法の一例を図２、図３を参照して説明する。
【００１９】
（イ）まず、図２の（ａ）に示すように、例えばホウケイ酸ガラスからなる基板１の表面を例えば３８０〜４００℃の硝酸カリウム溶液塩のようなイオン交換溶液に浸漬してカリウム、ナトリウム等の高屈折率元素をイオン交換することにより第１光導波路層２を形成する。その後、この第１光導波路層２上に例えば酸化チタン、酸化亜鉛、ニオブ酸リチウム、ＧａＡｓのような第１光導波路層２より屈折率の高い材料の層１１を、化学蒸着（ＣＶＤ）法等により形成する。
【００２０】
（ロ）次いで、図２の（ｂ）に示すように第１光導波路層２より屈折率の高い材料の層１１をフォトエッチング技術でパターニングすることにより第１光導波路層２の中央付近に外周が傾斜した形状の第２光導波路層４を形成する。
【００２１】
（ハ）つづいて、再度（イ）と同様に、全面に例えば酸化チタン、酸化亜鉛、ニオブ酸リチウム、ＧａＡｓのような第１光導波路層２より屈折率の高い材料の層を例えばＣＶＤ法等により形成する。その後、この層をフォトエッチング技術でパターニングすることにより図２の（ｃ）に示すように第１光導波路層２の両端部付近表面上に１対のグレーティング３を形成する。
【００２２】
（ニ）次いで、１対のグレーティング３および第２光導波路層４を含む第１光導波路層２上に、例えば感光性フッ素系樹脂のようなグレーティング３に比べて低屈折率を有する材料の被膜を塗布する。つづいて、この被膜に露光、現像処理を施すことにより図２の（ｄ）に示すように第２光導波路層４の表面に対応する箇所に矩形状の開口部６を有する感光性フッ素系樹脂からなる保護膜５を形成する。
【００２３】
（ホ）次に、図３の（ｅ）に示すように保護膜５の表面（開口部６の内面を除く）に例えば液晶表示装置に用いられるブラックマトリックスからなる迷光トラッピング層７をＣＶＤ法や真空蒸着により形成する。
【００２４】
（ヘ）さらに、図３の（ｆ）に示すように保護膜５の開口部６から露出した第２光導波路層４表面部分に酵素および発色色素を有する機能層８を形成する。具体的には、
１）酵素および発色色素を溶媒の存在下で架橋性の高分子（例えば光可溶性のポリビニルアルコール）と混合し、この溶液を保護膜５の開口部６から露出した第２光導波路層４表面部分にインクジェットまたはスピンコートにより塗布した後、光照射により光架橋性のポリビニルアルコールを架橋する；
２）酵素を色素の発色機能を発現する分子構造を持つ脂質分子の溶液を保護膜５の開口部６から露出した第２光導波路層４表面部分にインクジェットまたはスピンコートにより塗布した後、乾燥させる
等の方法により酵素および発色色素を有する機能層８を形成する。
【００２５】
（ト）次いで、保護膜５の開口部６から露出した機能層８上に、例えば有機モノマー、光反応開始剤およびデカン酸メチルのような貧溶媒を含む溶液を塗布、乾燥し、光重合させた後、洗浄処理することにより図３の（ｇ）に示すように多孔質膜９を形成する。この後、電界（例えばパルス状の電界）が印加されるメッシュ状導電性薄膜１０を多孔質膜９上に配置することにより前述した図１に示す光導波路型グルコースセンサを製造する。
【００２６】
次に図４を用いて検体１００のグルコースの測定方法について説明する。図４に示すように、電源１０２の陰極をメッシュ状導電性薄膜１０に配線し、電源１０２の陽極に電極板１０１を配線する。そして図１に示す光導波路型グルコースセンサの酵素および発色試薬を含む機能層８に設けられた多孔質膜９側に検体１００、例えば人体の皮膚の一部を接触させ、さらに検体１００の別の部位に電極板１０１を接触させる。ここで、この多孔質膜９の上部に配置されたメッシュ状導電性薄膜１０と電極板１０１の間に電源１０２から所望の電界（例えばパルス状電界）を印加すると、検体１００からグルコースを含む体液が多孔質膜９を通して機能層８に効率よく抽出される、いわゆる微浸襲作用がなされる。この体液中のグルコースは、機能層８でグルコースオキシダーゼ（ＧＯＤ）等との酸化酵素反応により過酸化水素（Ｈ_２Ｏ_２）を発生させ、さらに過酸化水素からペルオキシダーゼ（ＰＯＤ）等との酸化還元酵素反応によりラジカル酸素原子（Ｏ^＊）が生成される。このラジカル酸素原子により発色試薬を発色させる。この反応を模式すると(2)、(3)、(4)のようになる：
グルコース ＋ 酸化酵素（ＧＯＤ等）→ Ｈ_２Ｏ_２ ・・・・・(2)
Ｈ_２Ｏ_２ ＋ 酸化還元酵素（ＰＯＤ等）→ Ｏ^＊ ・・・・・(3)
Ｏ^＊ ＋ 発色試薬 → 発色 ・・・・・(4)
このような状態で、図１に示すように光源２１（例えば波長６５０ｎｍの半導体レーザ）および受光素子２２をそれぞれグルコースセンサの基板１の裏面左側および右側に配置する。そして光源２１からレーザ光を偏光フィルタ２３を通してグルコースセンサの基板１裏面側に入射する。そのレーザ光は基板１を通してグレーティング３と第１光導波路層２の界面で屈折されて、その第１光導波路層２を伝播される。第１光導波路層２を伝播されるレーザ光は、この第１光導波路層２より高屈折率の第２光導波路層４との界面で２つのモード（ＴＭモード、ＴＥモード）に分割され、ＴＭモードレーザ光は第１光導波路層２を、ＴＥモードレーザ光は第２光導波路層４を伝播する。このとき、機能層８における発色試薬の発色に基づく変化（例えば吸光度変化）によりこの機能層８直下の第２光導波路層４を伝播する光の強度が変化する。第１光導波路層２と第２光導波路層４を伝播した光は、受光素子２２側端付近において第１光導波路層２と第２光導波路層４の界面で再び結合、干渉するため、第２光導波路層４を伝播する光の強度変化を増幅できる。その結果、機能層８における検体１００（人体の皮膚下）のグルコースと酵素の反応による発色試薬の発色に基づく第２光導波路層４を伝播する光の極微な変化も、偏光フィルタ２４を通して受光素子２２で検出することが可能になる。
【００２７】
したがって、図１に示す構造の第１の実施の形態の光導波路型グルコースセンサによれば、酵素および発色試薬を含む機能層８の上部にメッシュ状導電性薄膜１０を配置し、図４に示した電源配線により、メッシュ状導電性薄膜１０と電極板１０１の間に電源１０２から所望の電界（例えばパルス状電界）を印加することにより、検体１００からグルコース（例えば人体の皮膚下のグルコースを含む体液）を機能層８に効率よく抽出する、いわゆる微浸襲作用を図ることができる。さらに光導波路を第１光導波路層２と第２光導波路層４により構成することで、機能層８における検体１００（人体の皮膚下）から抽出されたグルコースと酵素の反応による発色試薬の発色に基づく第２光導波路層４を伝播する光の極微な変化を、第１光導波路層２と第２光導波路層４の界面で検出できるため、抽出された体液中の極微量のグルコースを高感度で分析することができる。
【００２８】
また、図１に示すように１対のグレーティング３を含む第１光導波路層２上に保護膜５を形成することによって、第１光導波路層２および１対のグレーティング３に外部圧力が直接加わるのを防止できる。このため、第１光導波路層２およびグレーティング３に外部圧力が直接加わることに伴う部材の屈折率変化によって、その第１光導波路層２を伝播する光が外部に漏れるのを防止できる。しかも、保護膜５を第１光導波路層２に比べて低屈折率材料で形成することによって、その第１光導波路層２を伝播する光を第１光導波路層２と保護膜５との界面で効果的に全反射させて第１光導波路層２内に封じ込めることができるため、第１光導波路層２から光が外部に漏れるのを防止できる。その結果、検体１００中の極微量のグルコースをより高感度で分析することが可能になる。
【００２９】
さらに、迷光トラッピング層７を保護膜５の表面（開口部６の内面を除く）に形成することによって、第１光導波路層２を伝播する光が保護膜５との界面から保護膜５側に漏れた場合、その漏光を迷光トラッピング層７でトラップすることができる。
【００３０】
すなわち、第１光導波路層２を伝播する光が保護膜５との界面から保護膜５側に漏れると、保護膜５表面と外界（空気）との屈折率の差により漏光は保護膜５表面で全反射して第２光導波路層４に迷光として入射されるため、前述した検体１００中のグルコースの検出感度を低下させる。これに対し、迷光トラッピング層７を保護膜５の表面に形成することによって、漏光が保護膜５表面で全反射することなく迷光トラッピング層７でトラップできるため、その漏光が第２光導波路層４に迷光として入射するのを防止でき、検体１００中のグルコースをより高感度で分析することが可能になる。
【００３１】
さらに、酵素および発色試薬を含む機能層８を多孔質膜９で覆うことによって、機能層８に対する検体１００から抽出される（例えば体液）中のグルコース以外の蛋白や血球等の不純物の影響を防ぐ。つまり機能層８における酵素とグルコースの酵素反応による発色試薬の発色反応に基づく変化によりこの機能層８直下の第２光導波路層４を伝播する光の強度変化に対して作用する外乱を低減することができ、検体１００から抽出された体液中の極微量のグルコースをより一層高感度で分析することが可能になる。
【００３２】
（第２の実施の形態）
図５に示す光導波路型グルコースセンサは、第２光導波路層４上に発色試薬固定化層１３を形成し、この発色試薬固定化層１３上に酵素固定化層１４を形成した構造を有する。つまり、図５に示す光導波路型グルコースセンサは、前述した第１の実施の形態の機能層８を発色試薬固定化層１３と酵素固定化層１４の２つの層の分離した構造を有する。
【００３３】
発色試薬固定化層１３は、第２光導波路層４表面に例えばＮ，Ｎ−ビス（２−ヒドロキシ−３−サルフォプロピル）トリジンジカリウム塩や、３，３’，５，５’−テトラメチルベンジリデンのような発色試薬を、例えばアミノアルキルトリメトキシシランのようなシランカップリング剤または光架橋ポリビニルアルコールのような架橋高分子を介して固定化することにより形成される。
【００３４】
酵素固定化層１４は、例えばグルコースオキシダーゼのほかに、ペルオキシダーゼやムタロターゼを、例えば下記式(5)に示す構造式の化合物のような脂質の膜で固定化することにより形成される。
【００３５】
【化２】

図５に示す構成の光導波路型グルコースセンサにおいて、図４に示した配線と同じように、メッシュ状導電性薄膜１０と電極板１０１を電源１０２に配線する。そして酵素固定化層１４上に設けられた多孔質膜９側に検体１００、例えば人体の皮膚の一部を接触させ、さらに検体１００の別の部位に電極板１０１を接触させる。ここで、この多孔質膜９の上部に配置されたメッシュ状導電性薄膜１０と電極板１０１の間に電源１０２から所望の電界（例えばパルス状電界）を印加すると、検体１００からグルコースを含む体液が多孔質膜９を通して酵素固定化層１４に効率よく抽出される、いわゆる微侵襲作用がなされる。この体液中のグルコースは、酵素固定化層１４中で第１の実施の形態で示した反応と同じ酵素反応を起こし、この酵素反応によるラジカル酸素原子により酵素固定化層１４直下の発色試薬固定化層１３の発色色素を発色させる。このような状態で、図５に示すように光源２１（例えば波長６５０ｎｍの半導体レーザ）および受光素子２２をそれぞれグルコースセンサの基板１の裏面左側および右側に配置する。そして、光源２１からレーザ光を偏光フィルタ２３を通してグルコースセンサの基板１裏面側に入射する。レーザ光は基板１を通してグレーティング３と第１光導波路層２の界面で屈折されてその第１光導波路層２を伝播される。第１光導波路層２を伝播されるレーザ光は、この第１光導波路層２より高屈折率の第２光導波路層４との界面で２つのモード（ＴＭモード、ＴＥモード）に分割され、ＴＭモードレーザ光は第１光導波路層２を、ＴＥモードレーザ光は第２光導波路層４を伝播する。このとき、発色試薬固定化層１３の発色に基づく変化（例えば吸光度変化）によりこの発色試薬固定化層１３直下の第２光導波路層４を伝播する光の強度が変化する。第１光導波路層２と第２光導波路層４を伝播した光は、受光素子２２側端付近において第１光導波路層２と第２光導波路層４の界面で再び結合、干渉するため、第２光導波路層４を伝播する光の強度変化を増幅できる。その結果、酵素固定化層１４における検体１００（人体の皮膚下）から抽出された体液中のグルコースと酵素の反応による発色試薬固定化層１３の発色に基づく、第２光導波路層４を伝播する光の極微な変化を偏光フィルタ２４を通して受光素子２２で検出することが可能になる。
【００３６】
したがって、図５に示す構造の第２の実施の形態に係る光導波路型グルコースセンサによれば、酵素固定化層１４の上部にメッシュ状導電性薄膜１０を配置し、図４に示した電源配線により、メッシュ状導電性薄膜１０と電極板１０１の間に電源１０２から所望の電界（例えばパルス状電界）を印加することで、検体１００からグルコース（例えば人体の皮膚下のグルコースを含む体液）を酵素固定化層１４に効率よく抽出する、いわゆる微侵襲作用を図ることができる。さらに光導波路を第１光導波路層２と第２光導波路層４により構成することで、酵素固定化層１４および発色試薬固定化層１３における検体１００（人体の皮膚下）から抽出されたグルコースと酵素の反応による発色に基づく第２光導波路層４を伝播する光の極微な変化を、第１光導波路層２と第２光導波路層４の界面で検出できるため、検体１００から抽出された体液中の極微量のグルコースを高感度で分析することができる。
【００３７】
（第３の実施の形態）
図６に示す光導波路型グルコースセンサは、１対のグレーティング３の間に位置する第１光導波路層２上にこの第１光導波路層２より高屈折率の導電性材料、例えば酸化錫（ＳｎＯ₂）、ＩＴＯから作られる第２光導波路層１５を形成し、かつこの第２光導波路層１５に所望の電界（例えばパルス状電界）を印加する構造を有する。
【００３８】
このような図６に示す構成の光導波路型グルコースセンサにおいて、図４に示した配線の変形として、電源１０２の陰極を第２光導波路層１５に配線し、電源１０２の陽極に電極板１０１を配線する。そして酵素および発色試薬を含む機能層８上に設けられた多孔質膜９側に検体１００、例えば人体の皮膚の一部を接触させ、さらに検体１００の別の部位に電極板１０１を接触させる。ここで、この多孔質膜９の後方に配置された導電性材料から作られる第２光導波路層１５と電極板１０１の間に電源１０２から所望の電界（例えばパルス状電界）を印加すると、人体の皮膚下のグルコースを含む体液が多孔質膜９を通して機能層８に効率よく抽出される、いわゆる微侵襲作用がなされる。この体液中のグルコースは、機能層８の酵素との間で酵素反応がなされ、この酵素反応により生成されるラジカル酸素原子により機能層８中の発色試薬を発色させる。このような状態で、図６に示すように光源２１（例えば波長６５０ｎｍの半導体レーザ）および受光素子２２をそれぞれグルコースセンサの基板１の裏面左側および右側に配置する。光源２１からレーザ光を偏光フィルタ２３を通してグルコースセンサの基板１裏面側に入射すると、そのレーザ光は基板１を通してグレーティング３と第１光導波路層２の界面で屈折されてその第１光導波路層２を伝播される。第１光導波路層２を伝播されるレーザ光は、この第１光導波路層２より高屈折率の第２光導波路層１５との界面で２つのモード（ＴＭモード、ＴＥモード）に分割され、ＴＭモードレーザ光は第１光導波路層２を、ＴＥモードレーザ光は第２光導波路層１５を伝播する。このとき、機能層８における発色試薬の発色に基づく変化（例えば吸光度変化）によりこの機能層８直下の第２光導波路層１５を伝播する光の強度が変化する。第１光導波路層２と第２光導波路層１５を伝播した光は、受光素子２２側端付近において第１光導波路層２と第２光導波路層１５の界面で再び結合、干渉するため、第２光導波路層１５を伝播する光の強度変化を増幅できる。その結果、機能層８における検体１００（人体の皮膚下）から抽出されたグルコースと酵素の反応による発色試薬の発色に基づく第２光導波路層１５を伝播する光の極微な変化を偏光フィルタ２４を通して受光素子２２で検出することが可能になる。
【００３９】
したがって、図６に示す構造の第３の実施の形態に係る光導波路型グルコースセンサによれば、酵素および発色試薬を含む機能層８の下方に配置された第２光導波路層１５を導電性材料から作り、前述した電源配線により、この第２光導波路層１５と電極板１０１の間に電源１０２から所望のパルス状電界を印加することにより検体１００からグルコース（例えば人体の皮膚下のグルコースを含む体液）を機能層８に効率よく抽出する、いわゆる微侵襲作用を図ることができる。さらに光導波路を第１光導波路層２と第２光導波路層１５により構成することで、機能層８における検体１００（人体の皮膚下）から抽出されたグルコースと酵素の反応による発色試薬の発色に基づく第２光導波路層１５を伝播する光の極微な変化を第１光導波路層２と第２光導波路層１５の界面で検出できるため、検体１００から抽出された体液中の極微量のグルコースを高感度で分析することができる。
【００４０】
（第４の実施の形態）
図７に示す光導波路型グルコースセンサは、１対のグレーティング３の間に位置する第１光導波路層２上にこの第１光導波路層２より高屈折率の導電性材料、例えばＳｎＯ2、ＩＴＯから作られる第２光導波路層１５を形成し、この第２光導波路層１５上に第２の実施の形態で説明したのと同様な発色試薬固定化層１３を形成し、この発色試薬固定化層１３上に第２の実施の形態で説明したのと同様な酵素固定化層１４を形成した構造を有する。つまり、図７に示す光導波路型グルコースセンサは前述した第３の実施の形態の機能層８を発色試薬固定化層１３と酵素固定化層１４の２つの層の分離した構造を有する。なお、第２光導波路層１５に所望の電界（例えばパルス状電界）が印加される。
【００４１】
図７に示す構成の光導波路型グルコースセンサにおいて、図４に示した配線の変形として、電源１０２の陰極を第２光導波路層１５に配線し、電源１０２の陽極に電極板１０１を配線する。そして酵素固定化層１４上に設けられた多孔質膜９側に検体１００、例えば人体の皮膚の一部を接触させ、さらに検体１００の別の部位に電極板１０１を接触させる。ここで、この多孔質膜９の下方に配置された導電性材料から作られる第２光導波路層１５と電極板１０１の間に電源１０２から所望の電界（例えばパルス状電界）を印加すると、人体の皮膚下のグルコースを含む体液が多孔質膜９を通して酵素固定化層１４に効率よく抽出される、いわゆる微侵襲作用がなされる。この体液中のグルコースは、酵素固定化層１４の酵素との間で酵素反応がなされ、この酵素反応により生成されるラジカル酸素原子により酵素固定化層１４直下の発色試薬固定化層１３の発色試薬を発色させる。このような状態で、図７に示すように光源２１（例えば波長６５０ｎｍの半導体レーザ）および受光素子２２をそれぞれグルコースセンサの基板１の裏面左側および右側に配置する。光源２１からレーザ光を偏光フィルタ２３を通してグルコースセンサの基板１裏面側に入射すると、そのレーザ光は基板１を通してグレーティング３と第１光導波路層２の界面で屈折されてその第１光導波路層２を伝播される。第１光導波路層２を伝播されるレーザ光は、この第１光導波路層２より高屈折率の第２光導波路層１５との界面で２つのモード（ＴＭモード、ＴＥモード）に分割され、ＴＭモードレーザ光は第１光導波路層２を、ＴＥモードレーザ光は第２光導波路層１５を伝播する。このとき、発色試薬固定化層１３の発色に基づく変化（例えば吸光度変化）によりこの発色試薬固定化層１３直下の第２光導波路層１５を伝播する光の強度が変化する。第１光導波路層２と第２光導波路層１５を伝播した光は、受光素子２２側端付近において第１光導波路層２と第２光導波路層１５の界面で再び結合、干渉するため、第２光導波路層１５を伝播する光の強度変化を増幅できる。その結果、酵素固定化層１４における検体１００（人体の皮膚下）から抽出されたグルコースと酵素の反応による発色試薬固定化層１３の発色試薬の発色に基づく第２光導波路層１５を伝播する光の極微な変化を偏光フィルタ２４を通して受光素子２２で検出することが可能になる。
【００４２】
したがって、図７に示す構造の第４の実施の形態に係る光導波路型グルコースセンサによれば、酵素固定化層１４の下方に配置された第２光導波路層１５を導電性材料から作り、前述した電源配線により、この第２光導波路層１５と電極板１０１の間に電源１０２から所望のパルス状電界を印加することにより検体１００のグルコース（例えば人体の皮膚下のグルコースを含む体液）を酵素固定化層１４に効率よく抽出する、いわゆる微侵襲作用を図ることができる。さらに光導波路を第１光導波路層２と第２光導波路層１５により構成して、酵素固定化層１４および発色試薬固定化層１３における検体１００（人体の皮膚下）から抽出されたグルコースと酵素の反応による発色に基づく第２光導波路層１５を伝播する光の極微な変化を第１光導波路層２と第２光導波路層１５の界面で検出できるため、検体１００から抽出された体液中の極微量のグルコースを高感度で分析することができる。
【００４３】
（第５の実施の形態）
図８に示す光導波路型グルコースセンサは、図６に示した第３の実施の形態に係る光導波路型グルコースセンサと異なり、第１光導波路層２上に形成する第２光導波路層４が、第１光導波路層２より高屈折率の材料から形成されればよく、特に導電性を必要とはしない。その他の構造は第３の実施の形態に係る光導波路型グルコースセンサと同様である。
【００４４】
図８に示す構成の光導波路型グルコースセンサにおいては、酵素および発色試薬を含む機能層８上に設けられた多孔質膜９に、血液や体液等のグルコースを含む溶液を滴下する。この溶液中のグルコースは、第１の実施の形態で示したように機能層８の酵素との間で酵素反応を起こし、この酵素反応により生成されるラジカル酸素原子により機能層８中の発色試薬を発色させる。このような状態で、図８に示すように光源２１（例えば波長６５０ｎｍの半導体レーザ）および受光素子２２をそれぞれグルコースセンサの基板１の裏面左側および右側に配置し、光源２１からレーザ光を偏光フィルタ２３を通してグルコースセンサの基板１裏面側に入射すると、そのレーザ光は基板１を通してグレーティング３と第１光導波路層２の界面で屈折されてその第１光導波路層２を伝播される。第１光導波路層２を伝播されるレーザ光は、この第１光導波路層２より高屈折率の第２光導波路層４との界面で２つのモード（ＴＭモード、ＴＥモード）に分割され、ＴＭモードレーザ光は第１光導波路層２を、ＴＥモードレーザ光は第２光導波路層４を伝播する。このとき、機能層８における発色試薬の発色に基づく変化（例えば吸光度変化）によりこの機能層８直下の第２光導波路層４を伝播する光の強度が変化する。第１光導波路層２と第２光導波路層４を伝播した光は、受光素子２２側端付近において第１光導波路層２と第２光導波路層４の界面で再び結合、干渉するため、第２光導波路層４を伝播する光の強度変化を増幅できる。その結果、機能層８における溶液中のグルコースと酵素の反応による発色試薬の発色に基づく第２光導波路層４を伝播する光の極微な変化を偏光フィルタ２４を通して受光素子２２で検出することが可能になる。
【００４５】
したがって、図８に示す構造の第５の実施の形態に係る光導波路型グルコースセンサによれば、光導波路を第１光導波路層２と第２光導波路層４により構成して、機能層８における溶液中のグルコースと酵素の反応による発色試薬の発色に基づく第２光導波路層４を伝播する光の極微な変化を第１光導波路層２と第２光導波路層４の界面で検出できるため、溶液中の極微量のグルコースを高感度で分析することができる。
【００４６】
（第６の実施の形態）
図９に示す光導波路型グルコースセンサは、図７に示した第４の実施の形態に係る光導波路型グルコースセンサと異なり、第１光導波路層２上に形成する第２光導波路層４が、この第１光導波路層２より高屈折率の材料から形成されればよく、特に導電性を必要とはしない。その他の構造は第４の実施の形態に係る光導波路型グルコースセンサと同様である。
【００４７】
図９に示す構成の光導波路型グルコースセンサにおいては、酵素を含む酵素固定化層１４上に設けられた多孔質膜９に、血液や体液等のグルコースを含む溶液を滴下する。この溶液中のグルコースは、酵素固定化層１４の酵素との間で酵素反応を起こし、この酵素反応により生成されるラジカル酸素原子により酵素固定化層１４直下の発色試薬固定化層１３の発色試薬を発色させる。このような状態で、図９に示すように光源２１（例えば波長６５０ｎｍの半導体レーザ）および受光素子２２をそれぞれグルコースセンサの基板１の裏面左側および右側に配置し、光源２１からレーザ光を偏光フィルタ２３を通してグルコースセンサの基板１裏面側に入射すると、そのレーザ光は基板１を通してグレーティング３と第１光導波路層２の界面で屈折されてその第１光導波路層２を伝播される。第１光導波路層２を伝播されるレーザ光は、この第１光導波路層２より高屈折率の第２光導波路層４との界面で２つのモード（ＴＭモード、ＴＥモード）に分割され、ＴＭモードレーザ光は第１光導波路層２を、ＴＥモードレーザ光は第２光導波路層４を伝播する。このとき、発色試薬固定化層１３の発色に基づく変化（例えば吸光度変化）によりこの発色試薬固定化層１３直下の第２光導波路層４を伝播する光の強度が変化する。第１光導波路層２と第２光導波路層４を伝播した光は、受光素子２２側端付近において第１光導波路層２と第２光導波路層４の界面で再び結合、干渉するため、第２光導波路層４を伝播する光の強度変化を増幅できる。その結果、酵素固定化層１４における溶液中のグルコースと酵素の反応による発色試薬固定化層１３の発色に基づく第２光導波路層４を伝播する光の極微な変化を偏光フィルタ２４を通して受光素子２２で検出することが可能になる。
【００４８】
したがって、図９に示す構造の第６の実施の形態に係る光導波路型グルコースセンサによれば、光導波路を第１光導波路層２と第２光導波路層４により構成して、酵素固定化層１４および発色試薬固定化層１３における溶液中のグルコースと酵素の反応による発色に基づく第２光導波路層４を伝播する光の極微な変化を第１光導波路層２と第２光導波路層４の界面で検出できるため、溶液中の極微量のグルコースを高感度で分析することができる。
【００４９】
（第７の実施の形態）
図１０に示す光導波路型グルコースセンサは、第２光導波路層４上に発色試薬固定化層１３を形成し、この発色試薬固定化層１３上に酵素触媒固定化層１６を形成した構造を有する。つまり、図１０に示す光導波路型グルコースセンサは前述した第１の実施の形態の機能層８を発色試薬固定化層１３と酵素触媒固定化層１６の２つの層の分離した構造を有する。
【００５０】
発色試薬固定化層１３は、第２光導波路層４表面に例えばＮ，Ｎ−ビス（２−ヒドロキシ−３−サルフォプロピル）トリジンジカリウム塩や、３，３’，５，５’−テトラメチルベンジリデンのような発色試薬を、例えばアミノアルキルトリメトキシシランのようなシランカップリング剤または光架橋ポリビニルアルコールのような架橋高分子を介して固定化することにより形成される。
【００５１】
酵素触媒固定化層１６は、グルコースオキシダーゼと白金（Ｐｔ）とを含水ゲルに混合することにより形成される。この酵素触媒固定化層１６と前述の酵素固定化層１４との違いは、ペルオキシダーゼ等の酵素によってラジカル酸素原子を生成していた反応を、金属触媒を使用して起こすように、ペルオキシダーゼ等の酵素に代わって白金を含むようにした点である。
【００５２】
図１０に示す構成の光導波路型グルコースセンサにおいて、図４に示した配線と同じように、メッシュ状導電性薄膜１０と電極板１０１を電源１０２に配線する。そして酵素触媒固定化層１６上に設けられた多孔質膜９側に検体１００、例えば人体の皮膚の一部を接触させ、さらに検体１００の別の部位に電極板１０１を接触させる。ここで、この多孔質膜９の上部に配置されたメッシュ状導電性薄膜１０と電極板１０１の間に電源１０２から所望の電界（例えばパルス状電界）を印加すると、人体の皮膚下のグルコースを含む体液が多孔質膜９を通して酵素触媒固定化層１６に効率よく抽出される、いわゆる微侵襲作用がなされる。この体液中のグルコースは、酵素触媒固定化層１６でグルコースオキシダーゼとの酵素反応により過酸化水素（Ｈ２Ｏ２）を発生させる。さらに過酸化水素から白金を触媒としてラジカル酸素原子（Ｏ＊）が生成される。このラジカル酸素原子により酵素触媒固定化層１６直下の発色試薬固定化層１３の発色試薬を発色させる。つまり第１の実施の形態にて示した反応模式(2)〜(4)のうち、(3)が下記の(6)に置き換わる：
Ｈ２Ｏ２ ＋ Ｐｔ → Ｏ＊ ・・・・・(6)
グルコース量を測定する方法は、第２の実施の形態と同様であり、発色試薬固定化層１３の発色試薬を発色による第２光導波路層４を伝播する光の極微な変化を偏光フィルタ２４を通して受光素子２２で検出する。
【００５３】
したがって、図１０に示す構造の第７の実施の形態に係る光導波路型グルコースセンサによれば、酵素触媒固定化層１６の上部にメッシュ状導電性薄膜１０を配置し、図４に示した電源配線により、メッシュ状導電性薄膜１０と電極板１０１の間に電源１０２から所望の電界（例えばパルス状電界）を印加することで、検体１００からグルコース（例えば人体の皮膚下のグルコースを含む体液）を酵素触媒固定化層１６に効率よく抽出する、いわゆる微侵襲作用を図ることができる。さらに光導波路を第１光導波路層２と第２光導波路層４により構成して、酵素触媒固定化層１６および発色試薬固定化層１３における検体１００（人体の皮膚下）から抽出されたグルコースと酵素および触媒の反応による発色に基づく第２光導波路層４を伝播する光の極微な変化を第１光導波路層２と第２光導波路層４の界面で検出できるため、検体１００から抽出された体液中の極微量のグルコースを高感度で分析することができる。
【００５４】
（第８の実施の形態）
図１１に示す光導波路型グルコースセンサは、１対のグレーティング３の間に位置する第１光導波路層２上にこの第１光導波路層２より高屈折率の導電性材料、例えばＳｎＯ₂、ＩＴＯから作られる第２光導波路層１５を形成し、この第２光導波路層１５上に第７の実施の形態で説明したのと同様な発色試薬固定化層１３を形成し、この発色試薬固定化層１３上に第７の実施の形態で説明したのと同様な酵素触媒固定化層１６を形成した構造を有する。つまり、図１１に示す光導波路型グルコースセンサは前述した第３の実施の形態の機能層８を発色試薬固定化層１３と酵素触媒固定化層１６の２つの層の分離した構造を有する。なお、第２光導波路層１５に所望の電界（例えばパルス状電界）が印加される。
【００５５】
図１１に示す構成の光導波路型グルコースセンサにおいて、図４に示した配線の変形として、電源１０２の陰極を第２光導波路層１５に配線し、電源１０２の陽極に電極板１０１を配線する。そして酵素触媒固定化層１６上に設けられた多孔質膜９側に検体１００、例えば人体の皮膚の一部を接触させ、さらに検体１００の別の部位に電極板１０１を接触させる。ここで、この多孔質膜９の下方に配置された導電性材料から作られる第２光導波路層１５と電極板１０１の間に電源１０２から所望の電界（例えばパルス状電界）を印加すると、人体の皮膚下のグルコースを含む体液が多孔質膜９を通して酵素触媒固定化層１６に効率よく抽出される、いわゆる微侵襲作用がなされる。この体液中のグルコースは、前述した第７の実施の形態と同様に酵素触媒固定化層１６のグルコースオキシダーゼとの酵素反応により過酸化水素（Ｈ_２Ｏ_２）を発生させ、さらに過酸化水素から白金を触媒としてラジカル酸素原子（Ｏ^＊）が生成される。このラジカル酸素原子により酵素触媒固定化層１６直下の発色試薬固定化層１３の発色試薬を発色させる。
【００５６】
グルコース量を測定する方法は、第４の実施の形態と同様であり、発色試薬固定化層１３の発色試薬を発色による第２光導波路層１５を伝播する光の極微な変化を偏光フィルタ２４を通して受光素子２２で検出する。
【００５７】
したがって、図１１に示す構造の第８の実施の形態に係る光導波路型グルコースセンサによれば、酵素触媒固定化層１６の下方に配置された第２光導波路層１５を導電性材料から作り、前述した電源配線により、この第２光導波路層１５と電極板１０１の間に電源１０２から所望のパルス状電界を印加することにより検体１００からグルコース（例えば人体の皮膚下のグルコースを含む体液）を酵素固定化層１４に効率よく抽出する、いわゆる微侵襲作用を図ることができる。さらに光導波路を第１光導波路層２と第２光導波路層１５により構成して、酵素触媒固定化層１６および発色試薬固定化層１３における検体１００（人体の皮膚下）から抽出した体液中のグルコースと酵素および触媒の反応による発色に基づく第２光導波路層１５を伝播する光の極微な変化を第１光導波路層２と第２光導波路層１５の界面で検出できるため、検体１００中の極微量のグルコースを高感度で分析することができる。
【００５８】
（第９の実施の形態）
図１２に示す光導波路型グルコースセンサは、図１１に示した第８の実施の形態に係る光導波路型グルコースセンサと異なり、第１光導波路層２上に形成する第２光導波路層４が、この第１光導波路層２より高屈折率の材料から形成されればよく、特に導電性を必要とはしない。その他の構造は第８の実施の形態に係る光導波路型グルコースセンサと同様である。
【００５９】
図１２に示す構成の光導波路型グルコースセンサにおいては、酵素触媒固定化層１６上に設けられた多孔質膜９に、血液や体液等のグルコースを含む溶液を滴下する。この溶液中のグルコースは、前述した第７の実施の形態と同様に酵素触媒固定化層１６のグルコースオキシダーゼとの酵素反応により過酸化水素（Ｈ_２Ｏ_２）を発生させる。さらに過酸化水素から白金を触媒としてラジカル酸素原子（Ｏ^＊）が生成される。このラジカル酸素原子により酵素触媒固定化層１６直下の発色試薬固定化層１３の発色試薬を発色させる。
【００６０】
グルコース量を測定する方法は、第４の実施の形態と同様であり、発色試薬固定化層１３の発色試薬を発色による第２光導波路層４を伝播する光の極微な変化を偏光フィルタ２４を通して受光素子２２で検出する。
【００６１】
したがって、図１２に示す構造の第９の実施の形態に係る光導波路型グルコースセンサによれば、光導波路を第１光導波路層２と第２光導波路層４により構成して、酵素触媒固定化層１６および発色試薬固定化層１３における溶液中のグルコースと酵素および触媒の反応による発色に基づく第２光導波路層４を伝播する光の極微な変化を第１光導波路層２と第２光導波路層４の界面で検出できるため、溶液中の極微量のグルコースを高感度で分析することができる。
【００６２】
ここで、第２の実施の形態ないし第９の実施の形態に係る光導波路型グルコースセンサにおいて、グレーティング３を含む第１光導波路層２上に保護膜５を形成することによって、第１の実施の形態で説明したように第１光導波路層２およびグレーティング３に外部圧力が直接加わるのを防止でき、検体１００中の極微量のグルコースをより高感度で分析することが可能になる。
【００６３】
さらに、第２の実施の形態ないし第９の実施の形態に係る光導波路型グルコースセンサにおいて、迷光トラッピング層７を保護膜５の表面（開口部６の内面を除く）に形成することによって、第１の実施の形態で説明したように第１光導波路層２を伝播する光が保護膜５との界面から保護膜５側に漏れた場合、その漏光を迷光トラッピング層７でトラップすることができる。
【００６４】
さらに、第２の実施の形態ないし第９の実施の形態に係る光導波路型グルコースセンサにおいて、機能層８、酵素固定化層１４または酵素触媒固定化層１６を多孔質膜９で覆うことによって、機能層８、酵素固定化層１４または酵素触媒固定化層１６に対する検体１００（例えば体液）や溶液中のグルコース以外の不純物、例えば蛋白や血球の影響、つまり機能層８、酵素固定化層１４または酵素触媒固定化層１６および発色試薬固定化層１３における酵素反応と発色反応に基づく、機能層８または発色試薬固定化層１３直下の第２光導波路層（４または１５）を伝播する光の強度変化に対して作用する外乱を低減でき、検体１００から抽出された体液や溶液中の極微量のグルコースをより一層高感度で分析することが可能になる。
【００６５】
本発明の実施の形態を説明するために各図において示した各層と膜の厚さ、また隣接する層または膜との位置関係は、あくまでも例示であって、本発明の機能を実現するために限定したものでない。よって、本発明の機能が実現可能な範囲において、種々の各層と膜の厚さや位置関係が考え得ることは言うまでもない。
【００６６】
【発明の効果】
以上詳述したように本発明によれば、検体１００から抽出される、例えば体液中の極微量のグルコースを高感度かつ高精度で分析することが可能な光導波路型グルコースセンサを提供することができる。
【図面の簡単な説明】
【図１】本発明の第１の実施の形態に係る光導波路型グルコースセンサを示す断面図である。
【図２】図１の光導波路型グルコースセンサの製造工程を示す行程断面図である（その１）。
【図３】図１の光導波路型グルコースセンサの製造工程を示す行程断面図である（その２）。
【図４】図１の光導波路型グルコースセンサを用いた光導波路型グルコース測定方法の図である。
【図５】本発明の第２の実施の形態に係る光導波路型グルコースセンサを示す断面図である。
【図６】本発明の第３の実施の形態に係る光導波路型グルコースセンサを示す断面図である。
【図７】本発明の第４の実施の形態に係る光導波路型グルコースセンサを示す断面図である。
【図８】本発明の第５の実施の形態に係る光導波路型グルコースセンサを示す断面図である。
【図９】本発明の第６の実施の形態に係る光導波路型グルコースセンサを示す断面図である。
【図１０】本発明の第７の実施の形態に係る光導波路型グルコースセンサを示す断面図である。
【図１１】本発明の第８の実施の形態に係る光導波路型グルコースセンサを示す断面図である。
【図１２】本発明の第９の実施の形態に係る光導波路型グルコースセンサを示す断面図である。
【符号の説明】
１…基板
２…第１光導波路層
３…グレーティング
４…第２光導波路層
５…保護膜
６…開口部
７…迷光トラッピング層
８…機能層
９…多孔質膜
１０…メッシュ状導電性薄膜
１１…第１光導波路層より屈折率の高い材料の層
１３…発色試薬固定化層
１４…酵素固定化層
１５…第２光導波路層
１６…酵素触媒固定化層
２１…光源
２２…受光素子
２３…偏光フィルタ
２４…偏光フィルタ
１００…検体
１０１…電極板
１０２…電源[0001]
BACKGROUND OF THE INVENTION
The present invention relates to a glucose sensor, and more particularly to an optical waveguide glucose sensor for measuring a glucose concentration in a solution such as blood.
[0002]
[Prior art]
For example, JP-A-9-61346 discloses a planar optical waveguide type glucose sensor. This glucose sensor forms a pair of gratings where light enters and exits the substrate surface, forms a single optical waveguide layer on the substrate surface located between the gratings, and further recognizes molecules on this optical waveguide layer. It has a structure in which a film having a function and an information conversion function is formed.
[0003]
In a glucose sensor with such a structure, light such as laser light is incident on the optical waveguide layer through the grating in a state in which biomolecules contained in blood extracted from the specimen are in contact with a film having a molecular recognition function and an information exchange function. The evanescent wave is generated, and the amount of change in the evanescent wave caused by the reaction with the biomolecule contained in the blood or the like by the film on the optical waveguide layer is detected by the light receiving element that receives the light emitted from the grating, Analyzes biomolecules contained in blood and the like.
[0004]
[Problems to be solved by the invention]
However, the conventional glucose sensor has a single optical waveguide layer, and there is a limit in sensitivity in detecting the amount of change in the evanescent wave generated here, and blood extracted from the specimen from the film structure on the optical waveguide layer, etc. There is a problem that it is not suitable for analysis of a trace amount of biomolecules contained in.
[0005]
An object of the present invention is to provide an optical waveguide glucose sensor capable of analyzing a very small amount of glucose extracted from a specimen, for example, in body fluid with good sensitivity and high accuracy.
[0006]
[Means for Solving the Problems]
An optical waveguide glucose sensor according to the present invention includes a substrate, a first optical waveguide layer formed on the surface of the substrate, a pair of gratings formed on surfaces near both ends of the first optical waveguide layer, A second optical waveguide layer formed on the first optical waveguide layer located between the gratings and having a refractive index higher than that of the first optical waveguide layer, and an enzyme and a coloring reagent formed on the second optical waveguide layer. And a functional layer including the functional layer.
[0007]
Another optical waveguide glucose sensor according to the present invention includes a substrate, a first optical waveguide layer formed on the surface of the substrate, a pair of gratings formed on the surfaces near both ends of the first optical waveguide layer, A second optical waveguide layer formed on the first optical waveguide layer located between the pair of gratings and having a higher refractive index than the first optical waveguide layer; and a coloring reagent formed on the second optical waveguide layer. A coloring reagent-immobilizing layer containing the enzyme, and an enzyme-immobilizing layer containing an enzyme formed on the coloring reagent-immobilizing layer.
[0008]
Still another optical waveguide glucose sensor according to the present invention includes a substrate, a first optical waveguide layer formed on the surface of the substrate, and a pair of gratings formed on surfaces near both ends of the first optical waveguide layer. The second optical waveguide layer formed on the first optical waveguide layer positioned between the pair of gratings and having a higher refractive index than the first optical waveguide layer, and the color formed on the second optical waveguide layer It comprises a coloring reagent immobilization layer containing a reagent, and an enzyme catalyst immobilization layer containing an enzyme and a catalyst formed on the coloring reagent immobilization layer.
[0010]
DETAILED DESCRIPTION OF THE INVENTION
Next, first to ninth embodiments of the present invention will be described with reference to the drawings. In the following description of the drawings, the same or similar parts are denoted by the same or similar reference numerals. However, it should be noted that the drawings are schematic.
[0011]
(First embodiment)
In the optical waveguide glucose sensor according to the first embodiment of the present invention, as shown in FIG. 1, a substrate 1 made of glass, for example, and a first optical waveguide layer 2 having a higher refractive index than the substrate 1 on the surface thereof. Is formed. The pair of gratings (diffraction gratings) 3 has a higher refractive index than that of the first optical waveguide layer 2 and is formed on the surfaces near both ends of the first optical waveguide layer 2. Further, the second optical waveguide layer 4 having an inclined outer periphery has a higher refractive index than the first optical waveguide layer 2 and is formed on the first optical waveguide layer 2 positioned between the pair of gratings 3. ing.
[0012]
The protective film 5 has a lower refractive index than the grating 3 and is formed on the first optical waveguide layer 2 including the pair of gratings 3. A rectangular opening 6 is provided in the protective film 5 corresponding to the upper surface of the second optical waveguide layer 4. For example, the stray light trapping layer 7 made of a black matrix (pigment-containing resist) used in a liquid crystal display device is formed on the surface of the protective film 5 (excluding the inner surface of the opening 6).
[0013]
Note that the refractive index of the substrate 1 is η ₁ , The refractive index of the first optical waveguide layer 2 is η ₂ The refractive index of a pair of gratings 3 is η _Three , The refractive index of the second optical waveguide layer 4 is η _Four And the refractive index of the protective film 5 is η _Five , Then the magnitude relationship of their refractive indices is η _Four ≧ η _Three > Η ₂ > Η ₁ > Η _Five It becomes.
[0014]
The functional layer 8 containing an enzyme and a coloring reagent is formed on the surface portion of the second optical waveguide layer 4 exposed from the opening 6. For example, the porous film 9 made of a hydrogel is formed on the functional layer 8 exposed from the opening 6. The mesh conductive thin film 10 to which an electric field (for example, a pulsed electric field) is applied is disposed on the porous film 9. The first optical waveguide layer 2 is formed, for example, by ion exchange of an element such as potassium or sodium with a glass component. The pair of gratings 3 includes, for example, titanium oxide, zinc oxide, lithium niobate, gallium arsenide (GaAs), indium tin oxide (ITO), polyimide, tantalum oxide (Ta ₂ O ₅ ). The second optical waveguide layer 4 is made of, for example, titanium oxide, zinc oxide, lithium niobate, GaAs, ITO, polyimide, Ta ₂ O ₅ Made by. The protective film 5 is made of, for example, a fluororesin. Here, as the enzyme contained in the functional layer 8, for example, glucose oxidase as an oxidase, peroxidase as an oxidoreductase, and mutarotase for converting α-D-glucose into β-D-glucose are used. Yes. On the other hand, examples of the coloring reagent contained in the functional layer 8 include N, N-bis (2-hydroxy-3-sulfopropyl) tolidine dipotassium salt, 3,3 ′, 5,5′-tetramethylbenzylidene, and the like. Can be used.
[0015]
As the functional layer 8, for example,
1) A structure in which an enzyme and a coloring dye are immobilized with a crosslinked polymer;
2) A structure in which an enzyme is immobilized with a lipid molecule having a molecular structure that expresses the coloring function of the dye.
Etc.
[0016]
Examples of the crosslinked polymer used in the functional layer 8 of 1) include a polymer containing a hydrogen bonding functional group such as photocrosslinkable polyvinyl alcohol. Examples of the lipid molecule having a molecular structure that expresses the coloring function of the dye used in the functional layer 8 of 2) include those having the structural formula shown in the following formula (1).
[0017]
[Chemical 1]

The mesh conductive thin film 10 is made of, for example, a titanium thin film.
[0018]
Next, an example of a method for manufacturing the above-described optical waveguide glucose sensor will be described with reference to FIGS.
[0019]
(A) First, as shown in FIG. 2A, the surface of the substrate 1 made of, for example, borosilicate glass is immersed in an ion exchange solution such as a potassium nitrate solution salt at 380 to 400 ° C. The first optical waveguide layer 2 is formed by ion exchange of the high refractive index elements. Thereafter, a layer 11 made of a material having a higher refractive index than that of the first optical waveguide layer 2 such as titanium oxide, zinc oxide, lithium niobate, or GaAs is formed on the first optical waveguide layer 2 by a chemical vapor deposition (CVD) method or the like. To form.
[0020]
(B) Next, as shown in FIG. 2B, the outer periphery of the first optical waveguide layer 2 is formed near the center of the first optical waveguide layer 2 by patterning the layer 11 having a higher refractive index than that of the first optical waveguide layer 2 by a photoetching technique. The second optical waveguide layer 4 having an inclined shape is formed.
[0021]
(C) Subsequently, as in (a) again, a layer of a material having a higher refractive index than the first optical waveguide layer 2 such as titanium oxide, zinc oxide, lithium niobate, GaAs, etc. is applied to the entire surface, for example, by CVD. To form. Thereafter, this layer is patterned by a photo-etching technique to form a pair of gratings 3 on the surfaces near both ends of the first optical waveguide layer 2 as shown in FIG.
[0022]
(D) Next, on the first optical waveguide layer 2 including the pair of gratings 3 and the second optical waveguide layer 4, for example, a coating of a material having a lower refractive index than the grating 3 such as a photosensitive fluorine-based resin. Apply. Subsequently, by exposing and developing this film, a photosensitive fluorine-based resin having a rectangular opening 6 at a position corresponding to the surface of the second optical waveguide layer 4 as shown in FIG. A protective film 5 is formed.
[0023]
(E) Next, as shown in FIG. 3E, a stray light trapping layer 7 made of, for example, a black matrix used in a liquid crystal display device is formed on the surface of the protective film 5 (excluding the inner surface of the opening 6) by a CVD method or the like. Formed by vacuum evaporation.
[0024]
(F) Furthermore, as shown in FIG. 3F, a functional layer 8 having an enzyme and a coloring dye is formed on the surface portion of the second optical waveguide layer 4 exposed from the opening 6 of the protective film 5. In particular,
1) The surface portion of the second optical waveguide layer 4 in which the enzyme and the coloring dye are mixed with a crosslinkable polymer (for example, photosoluble polyvinyl alcohol) in the presence of a solvent, and this solution is exposed from the opening 6 of the protective film 5. After coating by inkjet or spin coating, photocrosslinkable polyvinyl alcohol is crosslinked by light irradiation;
2) Apply a solution of a lipid molecule having a molecular structure that expresses the coloring function of the dye to the surface portion of the second optical waveguide layer 4 exposed from the opening 6 of the protective film 5 by inkjet or spin coating, and then dry.
The functional layer 8 having an enzyme and a coloring dye is formed by such a method.
[0025]
(G) Next, on the functional layer 8 exposed from the opening 6 of the protective film 5, a solution containing an organic monomer, a photoinitiator and a poor solvent such as methyl decanoate is applied, dried, and photopolymerized. Thereafter, the porous film 9 is formed by performing a cleaning process as shown in FIG. Thereafter, a mesh-like conductive thin film 10 to which an electric field (for example, a pulsed electric field) is applied is disposed on the porous film 9 to manufacture the optical waveguide glucose sensor shown in FIG.
[0026]
Next, a method for measuring glucose in the specimen 100 will be described with reference to FIG. As shown in FIG. 4, the cathode of the power source 102 is wired to the mesh-like conductive thin film 10, and the electrode plate 101 is wired to the anode of the power source 102. Then, a specimen 100, for example, a part of the human skin is brought into contact with the porous membrane 9 provided on the functional layer 8 containing the enzyme and the coloring reagent of the optical waveguide glucose sensor shown in FIG. The electrode plate 101 is brought into contact with the site. Here, when a desired electric field (for example, a pulsed electric field) is applied from the power source 102 between the mesh-like conductive thin film 10 disposed on the porous film 9 and the electrode plate 101, a bodily fluid containing glucose from the specimen 100. Is effectively extracted into the functional layer 8 through the porous film 9, so-called micro-invasion action is performed. The glucose in the body fluid is converted into hydrogen peroxide (H by the oxidase reaction with glucose oxidase (GOD) etc. in the functional layer 8. ₂ O ₂ ) And oxygen radicals (O) by oxidoreductase reaction from hydrogen peroxide with peroxidase (POD) and the like. ^* ) Is generated. The coloring reagent is colored by the radical oxygen atoms. This reaction can be modeled as (2), (3), (4):
Glucose + oxidase (GOD, etc.) → H ₂ O ₂ (2)
H ₂ O ₂ + Oxidoreductase (POD etc.) → O ^* (3)
O ^* + Color reagent → Color development (4)
In such a state, as shown in FIG. 1, the light source 21 (for example, a semiconductor laser having a wavelength of 650 nm) and the light receiving element 22 are disposed on the left side and the right side of the back surface of the substrate 1 of the glucose sensor, respectively. Laser light from the light source 21 is incident on the back side of the substrate 1 of the glucose sensor through the polarizing filter 23. The laser light is refracted through the substrate 1 at the interface between the grating 3 and the first optical waveguide layer 2 and propagates through the first optical waveguide layer 2. Laser light propagating through the first optical waveguide layer 2 is divided into two modes (TM mode and TE mode) at the interface with the second optical waveguide layer 4 having a higher refractive index than the first optical waveguide layer 2. The TM mode laser light propagates through the first optical waveguide layer 2, and the TE mode laser light propagates through the second optical waveguide layer 4. At this time, the intensity of light propagating through the second optical waveguide layer 4 immediately below the functional layer 8 is changed by a change (for example, a change in absorbance) based on the color of the coloring reagent in the functional layer 8. The light propagated through the first optical waveguide layer 2 and the second optical waveguide layer 4 is coupled and interferes again at the interface between the first optical waveguide layer 2 and the second optical waveguide layer 4 in the vicinity of the end on the light receiving element 22 side. 2 A change in intensity of light propagating through the optical waveguide layer 4 can be amplified. As a result, a slight change in the light propagating through the second optical waveguide layer 4 based on the coloration of the coloring reagent due to the reaction between glucose and the enzyme in the specimen 100 (under the skin of the human body) in the functional layer 8 is also detected through the polarizing filter 24. 22 can be detected.
[0027]
Therefore, according to the optical waveguide glucose sensor of the first embodiment having the structure shown in FIG. 1, the mesh-like conductive thin film 10 is arranged on the upper part of the functional layer 8 containing the enzyme and the coloring reagent, and is shown in FIG. By applying a desired electric field (for example, a pulsed electric field) from the power source 102 between the mesh-like conductive thin film 10 and the electrode plate 101 by using the power supply wiring, glucose (for example, glucose under the skin of the human body is included) from the specimen 100. It is possible to achieve a so-called micro-invasion effect that efficiently extracts (body fluid) into the functional layer 8. Further, by forming the optical waveguide by the first optical waveguide layer 2 and the second optical waveguide layer 4, it is possible to develop the coloring reagent by the reaction of glucose and enzyme extracted from the specimen 100 (under the human skin) in the functional layer 8. Since the minute change of the light propagating through the second optical waveguide layer 4 can be detected at the interface between the first optical waveguide layer 2 and the second optical waveguide layer 4, a very small amount of glucose in the extracted body fluid is highly sensitive. Can be analyzed.
[0028]
Further, as shown in FIG. 1, by forming the protective film 5 on the first optical waveguide layer 2 including the pair of gratings 3, external pressure is directly applied to the first optical waveguide layer 2 and the pair of gratings 3. Can be prevented. For this reason, it is possible to prevent light propagating through the first optical waveguide layer 2 from leaking to the outside due to a change in the refractive index of the member caused by external pressure being directly applied to the first optical waveguide layer 2 and the grating 3. In addition, the protective film 5 is formed of a material having a lower refractive index than that of the first optical waveguide layer 2, so that light propagating through the first optical waveguide layer 2 is transmitted to the interface between the first optical waveguide layer 2 and the protective film 5. Thus, the light can be effectively totally reflected and contained in the first optical waveguide layer 2, so that light can be prevented from leaking outside from the first optical waveguide layer 2. As a result, an extremely small amount of glucose in the specimen 100 can be analyzed with higher sensitivity.
[0029]
Further, by forming the stray light trapping layer 7 on the surface of the protective film 5 (excluding the inner surface of the opening 6), the light propagating through the first optical waveguide layer 2 moves from the interface with the protective film 5 to the protective film 5 side. In the case of leakage, the leakage light can be trapped by the stray light trapping layer 7.
[0030]
That is, when light propagating through the first optical waveguide layer 2 leaks from the interface with the protective film 5 to the protective film 5 side, the light leakage is caused by the difference in refractive index between the surface of the protective film 5 and the outside (air). Therefore, the detection sensitivity of glucose in the specimen 100 described above is lowered since it is totally reflected and incident on the second optical waveguide layer 4 as stray light. On the other hand, by forming the stray light trapping layer 7 on the surface of the protective film 5, the light leakage can be trapped by the stray light trapping layer 7 without being totally reflected on the surface of the protective film 5. Can be prevented from entering as stray light, and glucose in the specimen 100 can be analyzed with higher sensitivity.
[0031]
Further, by covering the functional layer 8 containing the enzyme and the coloring reagent with the porous membrane 9, the influence of impurities other than glucose and proteins such as blood cells in the sample extracted from the specimen 100 (for example, body fluid) against the functional layer 8 is prevented. . That is, the disturbance acting on the intensity change of the light propagating through the second optical waveguide layer 4 immediately below the functional layer 8 is reduced by the change based on the color reaction of the color reagent due to the enzyme reaction between the enzyme and glucose in the functional layer 8. Therefore, it is possible to analyze a very small amount of glucose in the body fluid extracted from the specimen 100 with higher sensitivity.
[0032]
(Second Embodiment)
The optical waveguide glucose sensor shown in FIG. 5 has a structure in which a coloring reagent immobilization layer 13 is formed on the second optical waveguide layer 4 and an enzyme immobilization layer 14 is formed on the coloring reagent immobilization layer 13. That is, the optical waveguide type glucose sensor shown in FIG. 5 has a structure in which the functional layer 8 of the first embodiment described above is separated into two layers, the coloring reagent immobilization layer 13 and the enzyme immobilization layer 14.
[0033]
For example, N, N-bis (2-hydroxy-3-sulfopropyl) tolidine dipotassium salt or 3,3 ′, 5,5′-tetramethyl is formed on the surface of the second optical waveguide layer 4. It is formed by immobilizing a coloring reagent such as benzylidene via, for example, a silane coupling agent such as aminoalkyltrimethoxysilane or a crosslinking polymer such as photocrosslinked polyvinyl alcohol.
[0034]
The enzyme immobilization layer 14 is formed, for example, by immobilizing peroxidase or mutarotase in addition to glucose oxidase, for example, with a lipid membrane such as a compound having the structural formula shown in the following formula (5).
[0035]
[Chemical formula 2]

In the optical waveguide glucose sensor having the configuration shown in FIG. 5, as in the case of the wiring shown in FIG. The mesh-like conductive thin film 10 and the electrode plate 101 are used as a power source 102. Wiring. Then, the specimen 100, for example, a part of the human skin is brought into contact with the porous membrane 9 provided on the enzyme immobilization layer 14, and the electrode plate 101 is brought into contact with another part of the specimen 100. Here, when a desired electric field (for example, a pulsed electric field) is applied from the power source 102 between the mesh-like conductive thin film 10 disposed on the porous film 9 and the electrode plate 101, a bodily fluid containing glucose from the specimen 100. Is extracted through the porous membrane 9 into the enzyme-immobilized layer 14 so that a so-called microinvasive action is performed. The glucose in the body fluid causes the same enzyme reaction as the reaction shown in the first embodiment in the enzyme immobilization layer 14, and the coloring reagent immobilization immediately below the enzyme immobilization layer 14 is caused by radical oxygen atoms generated by this enzyme reaction. The coloring dye of the layer 13 is colored. In this state, as shown in FIG. 5, the light source 21 (for example, a semiconductor laser having a wavelength of 650 nm) and the light receiving element 22 are arranged on the left side and the right side of the back surface of the substrate 1 of the glucose sensor, respectively. Then, laser light from the light source 21 is incident on the back side of the substrate 1 of the glucose sensor through the polarizing filter 23. The laser light is refracted through the substrate 1 at the interface between the grating 3 and the first optical waveguide layer 2 and propagates through the first optical waveguide layer 2. Laser light propagating through the first optical waveguide layer 2 is divided into two modes (TM mode and TE mode) at the interface with the second optical waveguide layer 4 having a higher refractive index than the first optical waveguide layer 2. The TM mode laser light propagates through the first optical waveguide layer 2, and the TE mode laser light propagates through the second optical waveguide layer 4. At this time, the intensity of light propagating through the second optical waveguide layer 4 immediately below the color reagent immobilization layer 13 changes due to a change (for example, a change in absorbance) based on the color development of the color reagent immobilization layer 13. The light propagated through the first optical waveguide layer 2 and the second optical waveguide layer 4 is coupled and interferes again at the interface between the first optical waveguide layer 2 and the second optical waveguide layer 4 in the vicinity of the end on the light receiving element 22 side. 2 A change in intensity of light propagating through the optical waveguide layer 4 can be amplified. As a result, it propagates through the second optical waveguide layer 4 based on the color development of the coloring reagent immobilization layer 13 by the reaction between glucose and the enzyme in the body fluid extracted from the specimen 100 (under the human skin) in the enzyme immobilization layer 14. It becomes possible to detect a minute change in light by the light receiving element 22 through the polarizing filter 24.
[0036]
Therefore, according to the optical waveguide glucose sensor according to the second embodiment having the structure shown in FIG. 5, the mesh-like conductive thin film 10 is disposed on the enzyme immobilization layer 14 and the power supply wiring shown in FIG. By applying a desired electric field (for example, a pulsed electric field) from the power source 102 between the mesh-like conductive thin film 10 and the electrode plate 101, glucose (for example, body fluid containing glucose under the skin of the human body) is obtained from the specimen 100. A so-called microinvasive action that efficiently extracts the enzyme-immobilized layer 14 can be achieved. Furthermore, by configuring the optical waveguide by the first optical waveguide layer 2 and the second optical waveguide layer 4, glucose extracted from the specimen 100 (under the skin of the human body) in the enzyme-immobilized layer 14 and the coloring reagent-immobilized layer 13; Since a minute change in the light propagating through the second optical waveguide layer 4 based on the color developed by the reaction of the enzyme can be detected at the interface between the first optical waveguide layer 2 and the second optical waveguide layer 4, the body fluid extracted from the specimen 100 A very small amount of glucose in the medium can be analyzed with high sensitivity.
[0037]
(Third embodiment)
An optical waveguide glucose sensor shown in FIG. 6 is formed on a first optical waveguide layer 2 positioned between a pair of gratings 3, such as a conductive material having a higher refractive index than the first optical waveguide layer 2, such as tin oxide (SnO). ₂ ), A second optical waveguide layer 15 made of ITO is formed, and a desired electric field (for example, a pulsed electric field) is applied to the second optical waveguide layer 15.
[0038]
In the optical waveguide glucose sensor having the configuration shown in FIG. 6, as a modification of the wiring shown in FIG. 4, the cathode of the power source 102 is wired to the second optical waveguide layer 15, and the electrode plate 101 is attached to the anode of the power source 102. Wiring. Then, the specimen 100, for example, a part of the human skin is brought into contact with the porous film 9 provided on the functional layer 8 containing the enzyme and the coloring reagent, and the electrode plate 101 is brought into contact with another part of the specimen 100. Here, when a desired electric field (for example, a pulsed electric field) is applied from the power source 102 between the second optical waveguide layer 15 made of a conductive material disposed behind the porous film 9 and the electrode plate 101, the human body The body fluid containing glucose under the skin is efficiently extracted into the functional layer 8 through the porous membrane 9, so-called microinvasive action is performed. The glucose in the body fluid undergoes an enzymatic reaction with the enzyme in the functional layer 8, and the coloring reagent in the functional layer 8 is colored by radical oxygen atoms generated by this enzymatic reaction. In this state, as shown in FIG. 6, the light source 21 (for example, a semiconductor laser having a wavelength of 650 nm) and the light receiving element 22 are arranged on the left side and the right side of the back surface of the substrate 1 of the glucose sensor, respectively. When laser light from the light source 21 enters the back side of the substrate 1 of the glucose sensor through the polarizing filter 23, the laser light is refracted through the substrate 1 at the interface between the grating 3 and the first optical waveguide layer 2, and the first optical waveguide layer 2. Is propagated. The laser light propagated through the first optical waveguide layer 2 is divided into two modes (TM mode and TE mode) at the interface with the second optical waveguide layer 15 having a higher refractive index than the first optical waveguide layer 2. The TM mode laser light propagates through the first optical waveguide layer 2, and the TE mode laser light propagates through the second optical waveguide layer 15. At this time, the intensity of light propagating through the second optical waveguide layer 15 immediately below the functional layer 8 changes due to a change (for example, a change in absorbance) based on the color of the coloring reagent in the functional layer 8. The light propagated through the first optical waveguide layer 2 and the second optical waveguide layer 15 is recombined and interferes at the interface between the first optical waveguide layer 2 and the second optical waveguide layer 15 in the vicinity of the light receiving element 22 side end. 2 The intensity change of light propagating through the optical waveguide layer 15 can be amplified. As a result, a minute change in the light propagating through the second optical waveguide layer 15 based on the color of the coloring reagent generated by the reaction between glucose extracted from the specimen 100 (under the human skin) in the functional layer 8 and the enzyme is transmitted through the polarizing filter 24. Detection by the light receiving element 22 becomes possible.
[0039]
Therefore, according to the optical waveguide glucose sensor according to the third embodiment having the structure shown in FIG. 6, the second optical waveguide layer 15 disposed below the functional layer 8 containing the enzyme and the coloring reagent is used as the conductive material. By applying a desired pulsed electric field from the power source 102 between the second optical waveguide layer 15 and the electrode plate 101 by the power supply wiring described above, glucose from the specimen 100 (for example, glucose under the skin of the human body is included). A so-called microinvasive action that efficiently extracts body fluid) into the functional layer 8 can be achieved. Further, by forming the optical waveguide by the first optical waveguide layer 2 and the second optical waveguide layer 15, it is possible to develop a coloring reagent by the reaction of glucose and enzyme extracted from the specimen 100 (under the human skin) in the functional layer 8. Since the minute change of the light propagating through the second optical waveguide layer 15 can be detected at the interface between the first optical waveguide layer 2 and the second optical waveguide layer 15, a very small amount of glucose in the body fluid extracted from the specimen 100 can be detected. It can be analyzed with high sensitivity.
[0040]
(Fourth embodiment)
The optical waveguide glucose sensor shown in FIG. 7 is made of a conductive material having a higher refractive index than that of the first optical waveguide layer 2 on the first optical waveguide layer 2 located between the pair of gratings 3, such as SnO2 and ITO. The produced second optical waveguide layer 15 is formed, and the coloring reagent immobilization layer 13 similar to that described in the second embodiment is formed on the second optical waveguide layer 15, and this coloring reagent immobilization layer is formed. 13 has a structure in which an enzyme immobilization layer 14 similar to that described in the second embodiment is formed. That is, the optical waveguide glucose sensor shown in FIG. 7 has a structure in which the functional layer 8 of the above-described third embodiment is separated into two layers, a coloring reagent immobilization layer 13 and an enzyme immobilization layer 14. A desired electric field (for example, a pulsed electric field) is applied to the second optical waveguide layer 15.
[0041]
In the optical waveguide glucose sensor having the configuration shown in FIG. 7, as a modification of the wiring shown in FIG. 4, the cathode of the power source 102 is wired to the second optical waveguide layer 15, and the electrode plate 101 is wired to the anode of the power source 102. Then, the specimen 100, for example, a part of the human skin is brought into contact with the porous membrane 9 provided on the enzyme immobilization layer 14, and the electrode plate 101 is brought into contact with another part of the specimen 100. Here, when a desired electric field (for example, a pulsed electric field) is applied from the power source 102 between the second optical waveguide layer 15 made of a conductive material disposed below the porous film 9 and the electrode plate 101, the human body A body fluid containing glucose under the skin is efficiently extracted into the enzyme-immobilized layer 14 through the porous membrane 9, so-called microinvasive action is performed. The glucose in the body fluid undergoes an enzyme reaction with the enzyme of the enzyme immobilization layer 14, and the color reagent of the color reagent immobilization layer 13 immediately below the enzyme immobilization layer 14 by radical oxygen atoms generated by this enzyme reaction. To develop color. In such a state, as shown in FIG. 7, the light source 21 (for example, a semiconductor laser having a wavelength of 650 nm) and the light receiving element 22 are arranged on the left side and the right side of the back surface of the substrate 1 of the glucose sensor, respectively. When laser light from the light source 21 enters the back side of the substrate 1 of the glucose sensor through the polarizing filter 23, the laser light is refracted through the substrate 1 at the interface between the grating 3 and the first optical waveguide layer 2, and the first optical waveguide layer 2. Is propagated. The laser light propagated through the first optical waveguide layer 2 is divided into two modes (TM mode and TE mode) at the interface with the second optical waveguide layer 15 having a higher refractive index than the first optical waveguide layer 2. The TM mode laser light propagates through the first optical waveguide layer 2, and the TE mode laser light propagates through the second optical waveguide layer 15. At this time, the intensity of light propagating through the second optical waveguide layer 15 immediately below the color reagent immobilization layer 13 changes due to a change (for example, a change in absorbance) based on the color development of the color reagent immobilization layer 13. The light propagated through the first optical waveguide layer 2 and the second optical waveguide layer 15 is recombined and interferes at the interface between the first optical waveguide layer 2 and the second optical waveguide layer 15 in the vicinity of the light receiving element 22 side end. 2 The intensity change of light propagating through the optical waveguide layer 15 can be amplified. As a result, the light propagating through the second optical waveguide layer 15 based on the coloration of the color reagent of the color reagent immobilization layer 13 by the reaction between glucose extracted from the specimen 100 (under the human skin) in the enzyme immobilization layer 14 and the enzyme. It is possible for the light receiving element 22 to detect a slight change in the above through the polarizing filter 24.
[0042]
Therefore, according to the optical waveguide glucose sensor according to the fourth embodiment having the structure shown in FIG. 7, the second optical waveguide layer 15 disposed below the enzyme-immobilized layer 14 is made of a conductive material, and By applying a desired pulsed electric field from the power source 102 between the second optical waveguide layer 15 and the electrode plate 101 by the power supply wiring thus prepared, the glucose of the specimen 100 (for example, body fluid containing glucose under the skin of the human body) is enzyme A so-called microinvasive action that efficiently extracts to the immobilization layer 14 can be achieved. Furthermore, the optical waveguide is constituted by the first optical waveguide layer 2 and the second optical waveguide layer 15, and glucose and enzyme extracted from the specimen 100 (under the human skin) in the enzyme immobilization layer 14 and the coloring reagent immobilization layer 13. Since the minute change of the light propagating through the second optical waveguide layer 15 based on the color developed by the reaction of the first optical waveguide layer 2 can be detected at the interface between the first optical waveguide layer 2 and the second optical waveguide layer 15, A trace amount of glucose can be analyzed with high sensitivity.
[0043]
(Fifth embodiment)
The optical waveguide glucose sensor shown in FIG. 8 is different from the optical waveguide glucose sensor according to the third embodiment shown in FIG. 6 in that the second optical waveguide layer 4 formed on the first optical waveguide layer 2 includes: What is necessary is just to form from the material of higher refractive index than the 1st optical waveguide layer 2, and especially electroconductivity is not required. Other structures are the same as those of the optical waveguide glucose sensor according to the third embodiment.
[0044]
In the optical waveguide glucose sensor having the configuration shown in FIG. 8, a solution containing glucose such as blood or body fluid is dropped onto the porous film 9 provided on the functional layer 8 containing an enzyme and a coloring reagent. As shown in the first embodiment, the glucose in the solution causes an enzyme reaction with the enzyme of the functional layer 8, and the coloring reagent in the functional layer 8 is generated by radical oxygen atoms generated by the enzyme reaction. To develop color. In this state, as shown in FIG. 8, the light source 21 (for example, a semiconductor laser having a wavelength of 650 nm) and the light receiving element 22 are arranged on the left side and the right side of the back surface of the substrate 1 of the glucose sensor, respectively. When the laser beam is incident on the back side of the substrate 1 of the glucose sensor through 23, the laser light is refracted at the interface between the grating 3 and the first optical waveguide layer 2 through the substrate 1 and propagates through the first optical waveguide layer 2. Laser light propagating through the first optical waveguide layer 2 is divided into two modes (TM mode and TE mode) at the interface with the second optical waveguide layer 4 having a higher refractive index than the first optical waveguide layer 2. The TM mode laser light propagates through the first optical waveguide layer 2, and the TE mode laser light propagates through the second optical waveguide layer 4. At this time, the intensity of light propagating through the second optical waveguide layer 4 immediately below the functional layer 8 is changed by a change (for example, a change in absorbance) based on the color of the coloring reagent in the functional layer 8. The light propagated through the first optical waveguide layer 2 and the second optical waveguide layer 4 is coupled and interferes again at the interface between the first optical waveguide layer 2 and the second optical waveguide layer 4 in the vicinity of the end on the light receiving element 22 side. 2 A change in intensity of light propagating through the optical waveguide layer 4 can be amplified. As a result, it is possible for the light receiving element 22 to detect a minute change in the light propagating through the second optical waveguide layer 4 based on the coloring of the coloring reagent due to the reaction between glucose in the solution and the enzyme in the functional layer 8. become.
[0045]
Therefore, according to the optical waveguide glucose sensor according to the fifth embodiment having the structure shown in FIG. 8, the optical waveguide is configured by the first optical waveguide layer 2 and the second optical waveguide layer 4. Since a very small change in the light propagating through the second optical waveguide layer 4 based on the coloration of the coloring reagent due to the reaction between glucose and the enzyme in the solution can be detected at the interface between the first optical waveguide layer 2 and the second optical waveguide layer 4, A very small amount of glucose in the solution can be analyzed with high sensitivity.
[0046]
(Sixth embodiment)
The optical waveguide glucose sensor shown in FIG. 9 differs from the optical waveguide glucose sensor according to the fourth embodiment shown in FIG. 7 in that the second optical waveguide layer 4 formed on the first optical waveguide layer 2 is What is necessary is just to form from the material of higher refractive index than this 1st optical waveguide layer 2, and electroconductivity is not required especially. Other structures are the same as those of the optical waveguide glucose sensor according to the fourth embodiment.
[0047]
In the optical waveguide glucose sensor having the configuration shown in FIG. 9, a solution containing glucose such as blood or body fluid is dropped onto the porous membrane 9 provided on the enzyme-immobilized layer 14 containing the enzyme. Glucose in this solution causes an enzyme reaction with the enzyme of the enzyme immobilization layer 14, and the color reagent of the color reagent immobilization layer 13 immediately below the enzyme immobilization layer 14 by radical oxygen atoms generated by this enzyme reaction. To develop color. In this state, as shown in FIG. 9, the light source 21 (for example, a semiconductor laser having a wavelength of 650 nm) and the light receiving element 22 are respectively disposed on the left side and the right side of the back surface of the substrate 1 of the glucose sensor. When the laser beam is incident on the back side of the substrate 1 of the glucose sensor through 23, the laser light is refracted at the interface between the grating 3 and the first optical waveguide layer 2 through the substrate 1 and propagates through the first optical waveguide layer 2. Laser light propagating through the first optical waveguide layer 2 is divided into two modes (TM mode and TE mode) at the interface with the second optical waveguide layer 4 having a higher refractive index than the first optical waveguide layer 2. The TM mode laser light propagates through the first optical waveguide layer 2, and the TE mode laser light propagates through the second optical waveguide layer 4. At this time, the intensity of light propagating through the second optical waveguide layer 4 immediately below the color reagent immobilization layer 13 changes due to a change (for example, a change in absorbance) based on the color development of the color reagent immobilization layer 13. The light propagated through the first optical waveguide layer 2 and the second optical waveguide layer 4 is coupled and interferes again at the interface between the first optical waveguide layer 2 and the second optical waveguide layer 4 in the vicinity of the end on the light receiving element 22 side. 2 A change in intensity of light propagating through the optical waveguide layer 4 can be amplified. As a result, the light receiving element 22 transmits the minute change of the light propagating through the second optical waveguide layer 4 based on the color development of the coloring reagent immobilization layer 13 due to the reaction between glucose in the solution and the enzyme in the enzyme immobilization layer 14 through the polarizing filter 24. It becomes possible to detect with.
[0048]
Therefore, according to the optical waveguide glucose sensor according to the sixth embodiment having the structure shown in FIG. 9, the optical waveguide is constituted by the first optical waveguide layer 2 and the second optical waveguide layer 4, and the enzyme-immobilized layer. 14 and the coloring reagent immobilization layer 13, the minute change of the light propagating through the second optical waveguide layer 4 based on the color development caused by the reaction between glucose in the solution and the enzyme is changed between the first optical waveguide layer 2 and the second optical waveguide layer 4. Since it can be detected at the interface, a very small amount of glucose in the solution can be analyzed with high sensitivity.
[0049]
(Seventh embodiment)
The optical waveguide glucose sensor shown in FIG. 10 has a structure in which a coloring reagent immobilization layer 13 is formed on the second optical waveguide layer 4 and an enzyme catalyst immobilization layer 16 is formed on the coloring reagent immobilization layer 13. . That is, the optical waveguide glucose sensor shown in FIG. 10 has a structure in which the functional layer 8 of the first embodiment described above is separated into two layers, the coloring reagent immobilization layer 13 and the enzyme catalyst immobilization layer 16.
[0050]
For example, N, N-bis (2-hydroxy-3-sulfopropyl) tolidine dipotassium salt or 3,3 ′, 5,5′-tetramethyl is formed on the surface of the second optical waveguide layer 4. It is formed by immobilizing a coloring reagent such as benzylidene via, for example, a silane coupling agent such as aminoalkyltrimethoxysilane or a crosslinking polymer such as photocrosslinked polyvinyl alcohol.
[0051]
The enzyme catalyst immobilization layer 16 is formed by mixing glucose oxidase and platinum (Pt) in a hydrogel. The difference between the enzyme-catalyzed immobilization layer 16 and the enzyme immobilization layer 14 described above is that an enzyme such as peroxidase is used to cause a reaction that generates radical oxygen atoms by an enzyme such as peroxidase to occur using a metal catalyst. Instead of platinum.
[0052]
In the optical waveguide glucose sensor having the configuration shown in FIG. 10, the same as the wiring shown in FIG. The mesh-like conductive thin film 10 and the electrode plate 101 are used as a power source 102. Wiring. Then, the specimen 100, for example, a part of the human skin is brought into contact with the porous membrane 9 provided on the enzyme catalyst immobilization layer 16, and the electrode plate 101 is brought into contact with another part of the specimen 100. Here, when a desired electric field (for example, a pulsed electric field) is applied from the power source 102 between the mesh-like conductive thin film 10 disposed on the porous film 9 and the electrode plate 101, glucose under the skin of the human body is converted. A so-called microinvasive action in which the contained body fluid is efficiently extracted to the enzyme catalyst immobilization layer 16 through the porous membrane 9 is performed. The glucose in the body fluid generates hydrogen peroxide (H 2 O 2) by an enzyme reaction with glucose oxidase in the enzyme catalyst immobilization layer 16. Furthermore, radical oxygen atoms (O *) are generated from hydrogen peroxide using platinum as a catalyst. The color reagent of the color reagent immobilization layer 13 immediately below the enzyme catalyst immobilization layer 16 is colored by this radical oxygen atom. That is, in the reaction schemes (2) to (4) shown in the first embodiment, (3) is replaced with the following (6):
H2O2 + Pt → O * (6)
The method for measuring the amount of glucose is the same as in the second embodiment, and a minute change in the light propagating through the second optical waveguide layer 4 due to the coloring of the coloring reagent of the coloring reagent immobilization layer 13 is transmitted through the polarizing filter 24. Detection is performed by the light receiving element 22.
[0053]
Therefore, according to the optical waveguide type glucose sensor according to the seventh embodiment having the structure shown in FIG. 10, the mesh-like conductive thin film 10 is disposed on the enzyme catalyst immobilization layer 16, and the power source shown in FIG. By applying a desired electric field (for example, pulsed electric field) from the power source 102 between the mesh-like conductive thin film 10 and the electrode plate 101 by wiring, glucose (for example, body fluid containing glucose under the skin of the human body) from the specimen 100 is applied. Can be efficiently extracted into the enzyme catalyst-immobilized layer 16 to achieve a so-called microinvasive action. Further, the optical waveguide is constituted by the first optical waveguide layer 2 and the second optical waveguide layer 4, and glucose extracted from the specimen 100 (under the human skin) in the enzyme catalyst-immobilized layer 16 and the coloring reagent-immobilized layer 13; Extracted from the specimen 100 because a minute change in the light propagating through the second optical waveguide layer 4 based on the color developed by the reaction between the enzyme and the catalyst can be detected at the interface between the first optical waveguide layer 2 and the second optical waveguide layer 4. A very small amount of glucose in a body fluid can be analyzed with high sensitivity.
[0054]
(Eighth embodiment)
An optical waveguide type glucose sensor shown in FIG. 11 is formed on a first optical waveguide layer 2 positioned between a pair of gratings 3 and has a higher refractive index than that of the first optical waveguide layer 2, for example, SnO. ₂ A second optical waveguide layer 15 made of ITO is formed, and a coloring reagent immobilization layer 13 similar to that described in the seventh embodiment is formed on the second optical waveguide layer 15, and this coloring reagent is formed. It has a structure in which an enzyme catalyst immobilization layer 16 similar to that described in the seventh embodiment is formed on the immobilization layer 13. That is, the optical waveguide glucose sensor shown in FIG. 11 has a structure in which the functional layer 8 of the above-described third embodiment is separated into two layers, a coloring reagent immobilization layer 13 and an enzyme catalyst immobilization layer 16. A desired electric field (for example, a pulsed electric field) is applied to the second optical waveguide layer 15.
[0055]
In the optical waveguide glucose sensor having the configuration shown in FIG. 11, as a modification of the wiring shown in FIG. 4, the cathode of the power source 102 is wired to the second optical waveguide layer 15, and the electrode plate 101 is wired to the anode of the power source 102. Then, the specimen 100, for example, a part of the human skin is brought into contact with the porous membrane 9 provided on the enzyme catalyst immobilization layer 16, and the electrode plate 101 is brought into contact with another part of the specimen 100. Here, when a desired electric field (for example, a pulsed electric field) is applied from the power source 102 between the second optical waveguide layer 15 made of a conductive material disposed below the porous film 9 and the electrode plate 101, the human body A body fluid containing glucose under the skin is efficiently extracted into the enzyme catalyst-immobilized layer 16 through the porous membrane 9, so-called microinvasive action is performed. Glucose in the body fluid is converted into hydrogen peroxide (H by an enzyme reaction with glucose oxidase in the enzyme catalyst immobilization layer 16 as in the seventh embodiment. ₂ O ₂ ) And radical oxygen atoms (O ^* ) Is generated. The color reagent of the color reagent immobilization layer 13 immediately below the enzyme catalyst immobilization layer 16 is colored by this radical oxygen atom.
[0056]
The method for measuring the amount of glucose is the same as in the fourth embodiment, and a minute change in the light propagating through the second optical waveguide layer 15 due to the coloring of the coloring reagent of the coloring reagent immobilization layer 13 is transmitted through the polarizing filter 24. Detection is performed by the light receiving element 22.
[0057]
Therefore, according to the optical waveguide glucose sensor according to the eighth embodiment having the structure shown in FIG. 11, the second optical waveguide layer 15 disposed below the enzyme catalyst immobilization layer 16 is made of a conductive material, By applying a desired pulsed electric field from the power supply 102 between the second optical waveguide layer 15 and the electrode plate 101 by the power supply wiring described above, glucose (for example, body fluid containing glucose under the skin of the human body) is obtained from the specimen 100. A so-called microinvasive action that efficiently extracts the enzyme-immobilized layer 14 can be achieved. Further, the optical waveguide is constituted by the first optical waveguide layer 2 and the second optical waveguide layer 15, and is extracted from the specimen 100 (under the human skin) in the enzyme catalyst-immobilized layer 16 and the coloring reagent-immobilized layer 13. Since a very small change in the light propagating through the second optical waveguide layer 15 based on the color developed by the reaction of glucose, enzyme and catalyst can be detected at the interface between the first optical waveguide layer 2 and the second optical waveguide layer 15, A trace amount of glucose can be analyzed with high sensitivity.
[0058]
(Ninth embodiment)
Unlike the optical waveguide glucose sensor according to the eighth embodiment shown in FIG. 11, the optical waveguide glucose sensor shown in FIG. 12 has a second optical waveguide layer 4 formed on the first optical waveguide layer 2. What is necessary is just to form from the material of higher refractive index than this 1st optical waveguide layer 2, and electroconductivity is not required especially. Other structures are the same as those of the optical waveguide glucose sensor according to the eighth embodiment.
[0059]
In the optical waveguide glucose sensor having the configuration shown in FIG. 12, a solution containing glucose, such as blood or body fluid, is dropped onto the porous membrane 9 provided on the enzyme catalyst immobilization layer 16. Glucose in the solution is converted into hydrogen peroxide (H by an enzyme reaction with glucose oxidase in the enzyme catalyst immobilization layer 16 as in the seventh embodiment. ₂ O ₂ ). Furthermore, radical oxygen atoms (O ^* ) Is generated. The color reagent of the color reagent immobilization layer 13 immediately below the enzyme catalyst immobilization layer 16 is colored by this radical oxygen atom.
[0060]
The method for measuring the amount of glucose is the same as in the fourth embodiment, and the minute change of the light propagating through the second optical waveguide layer 4 due to the color development of the color development reagent immobilization layer 13 is transmitted through the polarization filter 24. Detection is performed by the light receiving element 22.
[0061]
Therefore, according to the optical waveguide glucose sensor according to the ninth embodiment having the structure shown in FIG. 12, the optical waveguide is constituted by the first optical waveguide layer 2 and the second optical waveguide layer 4, and is immobilized on the enzyme catalyst. In the first optical waveguide layer 2 and the second optical waveguide, a slight change in the light propagating through the second optical waveguide layer 4 based on the color development caused by the reaction between glucose in the solution, the enzyme and the catalyst in the layer 16 and the coloring reagent immobilization layer 13 is shown. Since it can be detected at the interface of the layer 4, a very small amount of glucose in the solution can be analyzed with high sensitivity.
[0062]
Here, in the optical waveguide glucose sensor according to the second embodiment to the ninth embodiment, the protective film 5 is formed on the first optical waveguide layer 2 including the grating 3 to thereby perform the first implementation. As described in the embodiment, it is possible to prevent external pressure from being directly applied to the first optical waveguide layer 2 and the grating 3, and it is possible to analyze a trace amount of glucose in the specimen 100 with higher sensitivity.
[0063]
Furthermore, in the optical waveguide glucose sensor according to the second to ninth embodiments, the stray light trapping layer 7 is formed on the surface of the protective film 5 (excluding the inner surface of the opening 6). As described in the first embodiment, when light propagating through the first optical waveguide layer 2 leaks from the interface with the protective film 5 to the protective film 5 side, the leakage light can be trapped by the stray light trapping layer 7. .
[0064]
Furthermore, in the optical waveguide glucose sensor according to the second to ninth embodiments, by covering the functional layer 8, the enzyme immobilization layer 14 or the enzyme catalyst immobilization layer 16 with the porous membrane 9, The influence of impurities other than glucose in the specimen 100 (for example, body fluid) and solution, such as proteins and blood cells, on the functional layer 8, the enzyme immobilization layer 14 or the enzyme catalyst immobilization layer 16, that is, the functional layer 8, the enzyme immobilization layer 14 or Intensity of light propagating through the second optical waveguide layer (4 or 15) immediately below the functional layer 8 or the color reagent immobilization layer 13 based on the enzyme reaction and color reaction in the enzyme catalyst immobilization layer 16 and the color reagent immobilization layer 13 Disturbances acting on changes can be reduced, and extremely small amounts of glucose in body fluids and solutions extracted from the specimen 100 can be analyzed with higher sensitivity.
[0065]
The thicknesses of the layers and films shown in the drawings for explaining the embodiment of the present invention, and the positional relationship between adjacent layers or films are merely examples, and the functions of the present invention are realized. Not limited. Therefore, it goes without saying that various layers and film thicknesses and positional relationships can be considered within the range in which the function of the present invention can be realized.
[0066]
【The invention's effect】
As described above in detail, according to the present invention, it is possible to provide an optical waveguide glucose sensor capable of analyzing a very small amount of glucose extracted from the specimen 100, for example, in a body fluid with high sensitivity and high accuracy. it can.
[Brief description of the drawings]
FIG. 1 is a cross-sectional view showing an optical waveguide glucose sensor according to a first embodiment of the present invention.
FIG. 2 is a process cross-sectional view illustrating a manufacturing process of the optical waveguide glucose sensor of FIG. 1 (No. 1).
3 is a process cross-sectional view illustrating a manufacturing process of the optical waveguide glucose sensor of FIG. 1 (part 2);
4 is a diagram of an optical waveguide glucose measuring method using the optical waveguide glucose sensor of FIG. 1; FIG.
FIG. 5 is a cross-sectional view showing an optical waveguide glucose sensor according to a second embodiment of the present invention.
FIG. 6 is a sectional view showing an optical waveguide glucose sensor according to a third embodiment of the present invention.
FIG. 7 is a cross-sectional view showing an optical waveguide glucose sensor according to a fourth embodiment of the present invention.
FIG. 8 is a cross-sectional view showing an optical waveguide glucose sensor according to a fifth embodiment of the present invention.
FIG. 9 is a sectional view showing an optical waveguide glucose sensor according to a sixth embodiment of the present invention.
FIG. 10 is a sectional view showing an optical waveguide glucose sensor according to a seventh embodiment of the present invention.
FIG. 11 is a cross-sectional view showing an optical waveguide glucose sensor according to an eighth embodiment of the present invention.
FIG. 12 is a sectional view showing an optical waveguide glucose sensor according to a ninth embodiment of the present invention.
[Explanation of symbols]
1 ... Board
2 ... 1st optical waveguide layer
3 ... Grating
4. Second optical waveguide layer
5 ... Protective film
6 ... Opening
7 ... Stray light trapping layer
8 ... Functional layer
9 ... Porous membrane
10 ... Mesh-like conductive thin film
11: Layer of material having a higher refractive index than the first optical waveguide layer
13 ... Coloring reagent immobilization layer
14 ... Enzyme immobilization layer
15 ... Second optical waveguide layer
16 ... Enzyme catalyst immobilization layer
21 ... Light source
22. Light receiving element
23 ... Polarizing filter
24 ... Polarizing filter
100 ... Sample
101 ... Electrode plate
102 ... Power supply

Claims (10)

A substrate,
A first optical waveguide layer formed on the substrate surface;
A pair of gratings respectively formed on surfaces near both ends of the first optical waveguide layer;
A second optical waveguide layer formed on the first optical waveguide layer located between the pair of gratings and having a higher refractive index than the first optical waveguide layer;
A functional layer containing an enzyme and a coloring reagent formed on the second optical waveguide layer;
An optical waveguide glucose sensor comprising: an electrode plate that can contact a specimen disposed on the functional layer . The optical waveguide glucose sensor according to claim 1, further comprising a power source that applies an electric field between the functional layer and the electrode plate.   The optical waveguide glucose sensor according to claim 1 or 2, further comprising a mesh-like conductive thin film on the functional layer.   The optical waveguide glucose sensor according to any one of claims 1 to 3, wherein the second optical waveguide layer is made of a conductive material.   The optical waveguide glucose sensor according to any one of claims 1 to 4, wherein the functional layer is a layer in which the enzyme and the coloring dye are immobilized with a crosslinked polymer.   The optical waveguide glucose sensor according to claim 5, wherein the crosslinked polymer contains a hydrogen bonding functional group.   The optical waveguide glucose sensor according to claim 1, wherein a porous membrane is further formed on the functional layer. A substrate,
A first optical waveguide layer formed on the substrate surface;
A pair of gratings respectively formed on surfaces near both ends of the first optical waveguide layer;
A second optical waveguide layer formed on the first optical waveguide layer located between the pair of gratings and having a higher refractive index than the first optical waveguide layer;
A coloring reagent immobilization layer containing a coloring reagent formed on the second optical waveguide layer;
An enzyme immobilization layer containing an enzyme formed on the color reagent immobilization layer;
An optical waveguide glucose sensor comprising a mesh-like conductive thin film provided on the enzyme immobilization layer . A substrate,
A first optical waveguide layer formed on the substrate surface;
A pair of gratings respectively formed on surfaces near both ends of the first optical waveguide layer;
A second optical waveguide layer formed on the first optical waveguide layer located between the pair of gratings and having a higher refractive index than the first optical waveguide layer;
A coloring reagent immobilization layer containing a coloring reagent formed on the second optical waveguide layer;
An enzyme catalyst immobilization layer comprising an enzyme and a catalyst formed on the color reagent immobilization layer;
An optical waveguide glucose sensor comprising a mesh-like conductive thin film provided on the enzyme immobilization layer .   The optical waveguide glucose sensor according to claim 1, wherein the enzyme contains at least glucose oxidase.

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