201200873 六、發明說明： [發明所屬技術領域] 本發明大致係有關於使用抗體(antibody)作為探針(probe)而利用電性感測 (electrosesning)所進行之偵檢(detection)及量測(measurement)。特定而言，本發明係 有關於可用於電性感測抗體探針偵檢及量測之感測器。 [先前技術] 在醫學及其相關應用領域中，使用生物晶片(biochip)來對檢測樣本中的目標物質 (target substrate)進行偵檢已屬習知。在諸如精確度及成本等因素考量之下，生物晶片 感測器(biochip sensor,或，biosensor)己被應用於特定目標物質之偵檢。若技術可行 且成本可以接受，對於任何可以想像到的用途而言，在檢測目標物質之出現與否之外， 若能更進一步進行定量量測，則生物晶片感測器顯然是更有其用處。例如，在生物醫療 (biomedical)應用之中，若能以，例如，由1至1〇，1至1〇〇,甚或更高的解析度，並 且維持精確度，而量測指出目標物質出現於樣本中之程度尺度，對於其原定感測應用目 的而言，顯然極具資訊性。 以光學感測(optical sensing)技術為基礎的生物晶片，是為今日全球生物感測技術所 常見者。此類晶片所依賴的是，需要利用龐大笨重且價格昂貴的精密光學儀器，進行光 學感測才得以讀取晶片上的檢測反應結果。為避開此些問題，利用電性感測原理的生物 晶片顯然是微小化與低成本的合理作法。在電性感測晶片的技術之中，當晶片置身於檢 測樣本中之後，其檢查(或者，感測)是電學性的檢查。由待檢測樣本上所感測到的可以 是阻抗值(impedance)，電容值（capacitance)，電阻值(resistance)，導電度' (conductance)，電流值(current),或任何其他有用的任何電性參數。 然而，直至目前為止，電性生物感測的技術有其限制，此係因絕大多數液體生物檢 測樣本，其本質皆未具電傳導性之故。圖3A及3B說明習知之電性感測技術，如何不 適於利用抗體探針進行抗原(antigen)檢測的情形。例如，圖3A以示意圖說明一種習知 技術感測晶片300所使用的是，例如固定在晶片正極312及負極314兩電極表面上的抗 體分子免疫球蛋白G (immunoglobulin G) 322,其中之電極可為，例如，金(Au)，銀 (Ag)，銅(Cu)或鎳(Ni)等之金屬薄膜。若要擁有實際用途，此系統必須要能夠容許量測 得到圖中大致以參考標號305所標示之環境，即感測晶片300的兩電極之間，其中電流 之變化量。 201200873 不過，在待檢測樣本導入之後，當其中的抗原分子332 (其大部份本質上皆為非導 體或不良導體)與固定(immobilized)在電極表面上的抗體分子322相互作用(interaction) 之後，此感測系統之電極之間的電傳導性，實際上仍是極為不良，如圖3B之示意圖所 示。因此，習知技藝之電性感測晶片技術，目前為止，僅適用於以酵素(enzyme)或催化 劑(catalyst)作為探針，以進行氧化還原作用產生電流電壓變化的生物晶片，其應用用途 因此極度受限。 [發明内容] 因此，本發明之目的係在於提供可利用電性感測抗體晶片偵檢及量測以供檢測各種 目標物質之存在。 因此，本發明之目的在於提供電性感測抗體探針偵檢及量測不同層度表現量以供檢 測各種目標物質之存在及各種目標物質之多寡。 另外，本發明之另一目的為提供一簡單，小型，低成本且可行之電性感測抗體探針 偵檢及量測方法，因此應用本方法之時，無須使用大型，高精準度以及昂貴的硬體設備。 此外，本發明之目的在於提供一電性感測抗體探針偵檢及量測方法，使其可用來檢 測各種不同的目標物質，並可廣泛應用在生物醫學之外的領域例如環境控制和工業界。 為達成上述其他目的本發明提供一種可供電性感測抗體探針偵檢及量測之硫醇基 寡噻吩(mercapto-oligothiophene)導電性誘昇化合物。電性感測之感測器包含有二電極 以及固定於至少一電極表面上的一層抗體。導電性誘昇分子被繋接於固定有該抗體之該 些電極上方及/或分佈於其間以増進兩電極間之電性傳導特性。就一特定層面而言，本 發明之電性感測晶片及其相關方法中之抗體探針分子可以說是「穿上了一件具電傳導性 的緊身衣」，此使得系統中的電傳導性實質上變得被「放大」到應用今日之儀器足以進 行精確感測的程度。電性感測晶片的諸如電阻值等的可量測電性參數，如此不但變得可 以偵檢並且得以判讀大小，其因此而變成可供進行解讀的有意義參數。 本發明提供一個電性感測器，實現了上述及其他目標待測物抗原之感測。此電性感 測器包括兩個分離且未連結的電極，以及固定於至少一電極表面上的一層抗體，並可與 高專一性抗原結合反應。當待測樣本之混合緩衝液流經晶片表面時，抗體接觸並結合抗 原，因而改變原先電極之導電特性，即在兩電極間顯示具電學感測所代表的訊號改變差 異量。 本發明的另一項具體證據，即提供一個電性感測器，實現了上述及其他目標待測物 201200873 抗原之感測◊此電性感測器包括兩個分離且未連結的電極，以及固定於至少一電極表面 上的一層抗體，並可與高專一性抗原結合反應。透過導電性誘昇分子促進抗體電極的導 電性，且改善兩個電極間電導度。當待測樣本之混合緩衝液流經晶片表面時，抗體接觸 並結合抗原，因而改變原先電極之導電特性，即在兩電極間顯示具電學感測所代表的訊 號改變差異量。 [圖式簡單說明] 圖1顯示一基本電性感測系統之架構。 圖2A及2B顯示電性感測晶片之兩種可能組構。 圖3A及3B說明習知技術電性感測如何不適於利用抗體探針進行抗原之檢測。 圖4A-4C分別顯示依據本發明感測晶片之一較佳實施例其製備及其對一様本進行感測 之情形。 圖5A-5C分別顯示依據本發明感測晶片之另一較佳實施例其製備及其對一様本進行感 測之情形。 圖6解釋本發明電性感測晶片及方法何以具有實質用途。 圖7A及7B本發明感測晶片進行表面修飾程序之兩實例，其可以提昇整個檢測系統中 的電傳導性。 圖8-10分別顯示利用本發明電性感測器進行數種樣本檢測之結果。 圖11說明本發明電性感測晶片一實施例之實體構造。 圖12顯示依據本發明一實施例之電性感測晶片處理機之流路組構。 [實施方式] 本發明利用提昇感測晶片系統(晶片環境以及與其發生反應之抗體分子本身)之電傳 導性，而得以做到有實質用途之電感測。特定而言，本發明之感測晶片及其相關方法中 之抗體探針分子可以說是「穿上了一件具有電傳導性的緊身衣」，此使得系統中的電傳 導性實質上變得被「放大j到應用今日之儀器足以進行精確感測的程度。感測晶片的諸 如電阻值等的可量測電性參數，如此不但變得可以偵檢並且得以判讀大小，其因此而變 成可供進行解讀的有意義參數。 依據本發明，被固定在感測晶片上作為檢測探針的抗體，實質上是由其非導電體本 質被轉變成為半導電性甚至是良好導電性的物質。此可容許被檢測樣本液體中目標物質 201200873 (在其與感測晶片上的抗體接觸並發生反應之後)的電性阻抗變化數值，不但變成可以被 儀器檢測出來，更能以足夠的精確度加以量測判別。量測讀取所得數值因此便可應用於 原定目標物質檢測用途中之解讀。 事實上，如同習於本技藝者所可理解，除了阻抗值以外，諸如系統的電容值等其他 電氣參數，亦皆由於本發明將整個系統的電傳導性加以提昇，而全皆變得可加以量測。 此外，除了解讀作為電阻值的倒數的嚴謹定義之外，「電導度」一詞在本發明之範疇内 亦可更為廣泛地被解讀為其系統中的電性傳導狀態。因此，「導電性誘昇」一詞在此應 被解讀為「電性傳導狀態之增進j。 因此，本發明之感測器及方法便能夠建立一種電性傳導之環境，其可以容許因為被 捕捉到的目標物質之出現於環境中，所導致之電傳導性的變動，變成不但可以被偵測 至IJ，並且可加以量測判讀。由於本發明之感測器及方法，實質上乃是將整個檢測樣本系 統的電特性偵檢範圍加以放大，因此其中電性質的任何變動，不論是阻抗> 電流或電容， 不論是以直流(DC)或任何選定頻率的交流(AC)加以量測，便皆很容易地可偵測到，並 可精確加以量化。其中之變化量因此即可成為被檢測樣本中目標物質存在量的一種量化 指標。 圖1顯示本發明一基本電性感測系統之架構建構在一片基材110上的感測晶片100 具有固定在其正電極112及負電極114表面上的整層的抗體探針12G其中的電極可為 例如，Au, Ag，Cu或Ni等之金屬薄膜。電極112及114係作為固定為特定抗原所選 定之抗體探針的實質基底。 以電性感測晶片100為基礎的本發明新穎電性感測技術之一種實施例系統，可與一 檢測電路與流體系統結合而提供一感測腔體(chamber) 102。在此腔體之中，檢測樣本與 晶片發生接觸，以容許懸浮在液態樣本中的目標抗原分子134得以變成被抗體探針120 所捕捉住的抗原132。 如同以下所將詳細說明者，圖1中之系統容許對檢測樣本中的目標抗原濃度進行精 確量測。此係利用在感測晶片的電極之間施加一電壓，並透過一電流量測儀器而進行量 測的，如同圖中所顯示者。 圖2A及2B顯示依據本發明一較佳實施例電性感測晶片之兩種可能組構 (configuration)»圖2A之感測晶片200A係採用典型平片形晶片形態其感測電極212A 及214A係在其基材210A上併排安置。此種平片形之晶片組構，係依賴其對應晶片讀 取處理裝置之配合，以形成其感測可以進行的一個樣本腔體。 201200873 相較之下，圖2B之電性感測晶片200B則具有管狀之組構(tubular configuration)， 其兩感測電極212B及214B係被設置在其管狀「基材j 210B内部表面的互相面對位置 上。利用此種管狀構形，只要在晶片被插入其對應之處理讀取機器時將其兩端加以封 閉，感測晶片200B便很容易地獲得一個樣本腔體202B。 圖4A-4C分別顯示依據本發明感測晶片之一較佳實施例其製備及其對一様本進行 感測之情形。注意到圖中所顯示電極，抗體，抗原及導電性誘昇分子等並未依正確尺度 比例繪示。為利於本發明之解釋說明，圖中所顯示者有部份係以誇大比例繪示。 圖4A顯示本發明一感測晶片之基本系統，在一個具有電傳導性的環境之中，其整 體導電性質,係利用導電性誘昇分子而得以提昇。在一較佳實施例中，薄膜形態的Au被 用來在感測晶片400的基材410上形成基本的正及負電極412與414。諸如Ag、Cu及 Ni等的金屬，於本發明中亦得以被利用來製作電極。依用途之不同，適當的合金，例 如，銦錫氧化物(indium tin oxide, ITO)，亦可使用。 具電傳導性之分子442被鍵結在電極上，此係如圖中所繪示的，其係被固定在電極 表面上。依據本發明，此些分子因而形成了固定在電極表面的導電性誘昇分子。當本發 明之晶片被使用時，此可令本發明之基本感測系統，得以提供一個增強的電性傳導環 境，而這是由於該些導電性誘昇分子修飾了感測晶片的表面性質，其結果使得裸感測晶 片系統的電傳導性得以提昇。亦即，在抗體探針分子出現在晶片系統中之後，正電極與 負電極之間的導電性因此大為增進。由於圖中以405A所標示的，電極412及414之間 的大為改善的電性傳導環境之故，此時系統便可以在感測晶片400的電極之間感測抗原 存在前後的電流變化。 適於用作導電性誘昇的物質包括，但不限於，寡噻吩-矽烷(oligothiophene-silane)， 寡噻吩-硫醇(oligothiophene-thiol)，1-苯基寡噻吩((l-phenyl)-oligothiophene), 2-苯基 寡噻吩((2-phenyl)-oligothiophene)，支鏈寡噻吩(side-arm oligothiophene),寡苯基噻 吩寡聚物(oligophenyl oligothiophene)，以及其衍生物等。 在圖4B中，針對特定目標抗原之探針用途的抗體422,被加至感測晶片400的導 電性誘昇分子層上，並與其共價鍵鍵結。因為具有此層固定化抗體，在此一階段(即當 目標抗原尚未出現時)，感測晶片的導電值於此已具有電導性的環境405B中，會有些減 降，但仍是在易於使用儀器進行量測的範圍之中。 當抗體422出現之後，圖4B中的晶片400便形成了備妥可針對其特定目標抗原分 子進行電性感測應用的一只電抗體感測器(electric antibody sensor)。就任何預先設定的 201200873 感測用途而言，其對應之特定非導電性抗體分子需要先被固定在晶片上。例如，免疫球 蛋白G (immunoglobulin G)分子可以被使用作為檢測諸如S100，甲型胎兒蛋白 (alpha-fetoprotein)，以及心肌鈣蛋白I (tropolin I)等的抗體探針。系統的整體電傳導 性會有所減降，其程度反映了探針之出現在系統中的事實。電導度的此一變化值變成了 檢測時之量測參考基準值。 圖4C顯示利用將目標抗原曝露給固定在晶片上的探針抗體，而進行電性感測的情 況。圖48中已備妥探針可供進行感測的晶片400被曝露在一檢測樣本之下。由於針對 特定目標可進行檢測的探針抗體422己被鍵結固定在晶片上，存在於樣本之中的目標， 亦即，抗原432,便被抗體所捕捉，或者是說，與抗體產生相互作用(interaction)反應。 隨著被捕捉住的抗原分子432之出現在系統之中，整個電性傳導環境405C的整體 電導度便隨著進一步變動(即，相較於圖4B)，而其電性阻抗量測讀數之變動(即，作為 電極間之電流而被量測者)，便成為系統中所出現抗原數量之程度的一個指標。 依據本發明所進行之電性感測，當含有非導電性目標抗原的樣本，被導入圖4C之 感測晶片所提供的流體偵檢量測環境内時，系統的整體電傳導性便會隨之變動減降。此 種減降係以量測所得電流值的對應減降加以反映。其變動程度係與代表被晶片所捕捉住 目標抗原的數量成比例。不過，應予注意的是，在某些情況下，檢測樣本内某些抗原之 與抗體探針結合(binding)後，相較於該些抗原尚未出現在系統中之前，反有可能會導致 電導度的增升。 圖5A-5C分別顯示依據本發明感測晶片之另一較佳實施例，其製備及其對一様本 進行感測之情形。圖5A-5C所描述之實例與圖4A-4C中所顯示者，除了其感測晶片的 物理構造，係採用互相面對的電極設置定位以外，兩者實質上是相同的。依本發明之推 論(但本發明不應受限於此推論)，此種電極對置之組構，應可能因其相對於圖4A-4C之 平片形組構之較佳導電性質，而得以容許產生較佳的感測性能。 圖6之示意圖解釋本發明電性感測晶片及方法何以具有其實質用途。圖中之曲線顯 示一檢測樣本之導電性質，相較於樣本中所出現之抗原濃度的相對關係。 圖6中垂直軸，即電傳導性質軸，上的符號A，B，C，D，D'及D",分別係為感 測晶片在其製備過程中各個不同階段的電傳導性： A: 基材 B: 電極 C: 電導度誘昇 201200873 D, D', D〃： 抗體探針加入 為了要對樣本内所存在的目標抗原，進行一個寬廣範圍濃度的量測，習知技藝係在 圖中的小電流量測讀取範圍(BD^或BD",依所加入之探針會稍微減降或增升其整體電傳 導性而定)内試圖量測樣本的導電性讀數。其電流值的讀數範圍小到無法有實際用途， 不但難以判別目標物質有否存在，更無庸說能夠以可接受的讀取解析度得出樣本内的抗 原濃度曲線，E'或E"。 相較之下，若依本發明使用導電性誘昇分子，目標物質的檢測範圍(BD)，就某一層 面之意義而言，乃是被實質放大了，其因而可以容許以良好的解析度，換言之即較佳的 精準度，而讀取判定目標濃度。這是由於，如圖6中的特性曲線E所清楚顯現的，不論 液體樣本環境中目標物質濃度與在其中之對應測得電流之間是為線性或非線性關係，在 寬廣的量測對應範圍内所進行的目標偵檢，當然可使儀器讀數的解讀遠較為容易之故。 圖7A及7B係本發明感測晶片進行表面修飾程序之兩實例，其可以提昇整個檢測 系統中的電導度。圖8-10則分別顯示利用本發明感測器進行數種樣本檢測之實驗結果。 圖8顯示S100抗原於實驗中，在-0.2V的量測電壓之下，相對於抗-(S-100蛋白)單 株抗體(anti-SlOO monoclonal antibody)的結合曲線(binding curve)。利用抗-(S-100 蛋 白)單株抗體(anti-S100 monoclonal antibody)(mAb)作為探針的感測晶片，被用來檢測 S100抗原。圖8中之結合曲線係在-0.2V檢測電壓之下所獲得的，其顯現了一個清楚可 判讀的關係曲線，並將每毫升(milliliteiO 0至200微克(microgram)的一整個樣本中目標 物質濃度範圍，對應到0至50毫微安培(nano-ampere)的感測電流尺度上。牛血清蛋白 (BSA，bovine serum albumin)在此實驗中被使用作為控制組樣本。 圖9顯示甲型胎兒蛋白(alpha fetoprotein)於實驗中，在-0.5V的量測電壓之下，相 對於抗-(甲型胎兒蛋白)抗體（anti-alpha fetoprotein monoclonal antibody),於不同抗 原濃度下的結合曲線。利用抗-(甲型胎兒蛋白)抗體(anti-alpha fetoprotein monoclonal antibody)(mAb)作為探針的感測晶片，被用來檢測甲型胎兒蛋白抗原。圖中之結合曲線 係在-0.5V檢測電壓之下，以每毫升0至1微克(microgram)濃度範圍的檢測抗原所獲得 的。此實驗其顯現了一個清楚可判讀的關係曲線，其將樣本中目標物質濃度範圍，對應 到0至500毫微安培的感測電流尺度上。BSA在此實驗中被使用作為控制組樣本。 圖10顯示抗-(心肌釣蛋白I)單株抗體(anti-(tropolin I) monoclonal antibody)於實 驗中，在-0.2V的量測電壓之下，相對於心肌鈣蛋白I (tropolin)的結合曲線。利用抗-(心 肌錦蛋白I)單株抗體(anti-(tropolin I) monoclonal antibody)(mAb)作為探針的感測晶 201200873 片被用來檢測心肌鈣蛋白I (tropolin)抗原。圖中之結合曲線係在-0.2V檢測電壓之下， 以每毫升〇至200微克濃度範圍的檢測樣本所獲得的。此實驗其顯現了一個清楚可判讀 的關係曲線，其將樣本中目標物質濃度範圍，對應到0至30毫微安培的感測電流尺度 上。BSA在此實驗中被使用作為控制組樣本。 圖11顯示本發明感測晶片一實施例之實體構造。在此平片形晶片組構的實例之中， 具平面電極架構的感測晶片1100擁有玻璃基材Ilia其上鍍有電極1112 1114及11½ 以及與印刷電路板中常見者相類似的邊緣連接電極116Z Π64及1166»於此實例之中， 此些電極包含有正1112及負1114兩功能性質電極之間的一個中間參考電極1116>其中 的某些邊緣連接電極則可預留供未來新功能之使用。可供感測晶片之功能性電極應用的 典型金屬材質導電鍍層包括有Au, Ag, Cu，Ni等，或者，其亦可使用軟性材料電極 或其他具導電性的材料。可供參考電極使用的典型惰性金屬層包括有Pt以及其他。可 將晶片電性連接至一感測電路的邊緣，其連接電極可以採用PCB技術中所常見者。此 些片層的典型厚度約為2,000埃(angstrom)。 圖12顯示依據本發明一實施例之電抗體感測器之整個系統架構„此感測器實例中， 使用了兩條液體流路1272及1274。當感測晶片1200,其亦可以被選擇性地承載於一晶 片載台1208上者，被應用於指定目標物質(抗原)的檢測時，其樣本係被注入器1284經 由載入通路而載入至樣本迴路1282肉如圖中的虛線所顯承接著利用一部栗浦(pump) 1127,由儲槽1286所儲存的一緩衝液，將樣本泵入以通過感測晶片電極表面。檢測樣 本及緩衝液兩者接著即離開而被排入一廢液槽1288内以待後續處理。 應注意的是，對於本發明之新穎性重點而言，以上配合所附圖式所說明之各種較佳 實施例，僅只詳細描述了本發明於感測晶片系統中，施行導電性誘昇的數種可能作法之 其中一種，亦即，將導電性誘昇分子亦同時作為抗體對電極的連結物(linkage)，即圖4 及5中所描繪者。不過，如同習於本技藝者所可以理解的，導電性誘昇分子於本發明系 統中的其他應用安排，亦同樣是可能且可行的。例如，其中稍有不同的安排，包含了利 用導電性誘昇分子來修飾抗體，將導電性誘昇分子加入至流經感測晶片表面的緩衝液 中，或於樣本之中加入使用，甚或上述此些安排的任何組合，其如同己經實驗所證實者 全皆是屬可行的作法。 雖然以上之說明文字，應己利用特定實施例，而完整地描述了本發明，但其各種修 改，變化構造及其等效者仍屬可以實施者。因此，以上說明文字及相關圖式，即不應被 作為本發明後列專利範圍範疇之限定性解讀。 10201200873 VI. OBJECTS OF THE INVENTION: 1. Field of the Invention The present invention generally relates to detection and measurement using electrosensing using an antibody as a probe. ). In particular, the present invention relates to sensors that are useful for the detection and measurement of electrophysiological antibody probes. [Prior Art] In the field of medicine and related fields, it is conventional to use a biochip to detect a target substrate in a test sample. Biochip sensors (or biosensors) have been applied to the detection of specific target substances under considerations such as accuracy and cost. If the technology is feasible and the cost is acceptable, the biochip sensor is obviously more useful for any imaginable use, in addition to detecting the presence or absence of the target substance, if further quantitative measurement is possible. . For example, in biomedical applications, if, for example, 1 to 1 〇, 1 to 1 〇〇, or even higher resolution, and maintaining accuracy, the measurement indicates that the target substance appears The degree scale in the sample is obviously very informative for the purpose of the original sensing application. Biochips based on optical sensing technology are common to today's global biosensing technology. Such wafers rely on the need to utilize optical instruments that are bulky and expensive to perform optical sensing to read the results of the detection on the wafer. In order to avoid these problems, biochips utilizing the principle of electro-sensing are obviously a reasonable method of miniaturization and low cost. Among the techniques of electro-sensing wafers, after the wafer is placed in the test sample, its inspection (or sensing) is an electrical inspection. Sensed by the sample to be detected may be impedance, capacitance, resistance, conductivity, current, or any other useful electrical property. parameter. However, until now, the technique of electrical biosensing has its limitations, which is due to the fact that most liquid bioassay samples are not electrically conductive. Figures 3A and 3B illustrate how conventional electro-sensing techniques are not suitable for antigen detection using antibody probes. For example, FIG. 3A schematically illustrates a conventional technique for sensing a wafer 300 using, for example, an antibody molecule immunoglobulin G 322 immobilized on the surface of both the positive electrode 312 and the negative electrode 314 of the wafer, wherein the electrode can be For example, a metal film of gold (Au), silver (Ag), copper (Cu) or nickel (Ni). For practical use, the system must be capable of permitting an environment generally indicated by reference numeral 305 in the figure, i.e., sensing the amount of change in current between the two electrodes of wafer 300. 201200873 However, after the introduction of the sample to be tested, after the antigen molecule 332 therein (which is mostly a non-conductor or a poor conductor in nature) interacts with the antibody molecule 322 immobilized on the surface of the electrode, The electrical conductivity between the electrodes of the sensing system is still extremely poor, as shown in the schematic of Figure 3B. Therefore, the electro-sensing wafer technology of the prior art is only applicable to a biochip which uses an enzyme or a catalyst as a probe to perform a redox reaction to generate a current-voltage change, and its application is extremely extreme. Limited. SUMMARY OF THE INVENTION Accordingly, it is an object of the present invention to provide for the detection and measurement of electrical sensing antibody wafers for the detection of the presence of various target substances. Accordingly, it is an object of the present invention to provide an electrical sensing antibody probe for detecting and measuring different levels of expression for detecting the presence of various target substances and the amount of various target substances. In addition, another object of the present invention is to provide a simple, small, low-cost and feasible electro-detection antibody probe detection and measurement method, so that the method does not need to use large, high precision and expensive when applying the method. Hardware equipment. In addition, the object of the present invention is to provide an electrosensing antibody probe detection and measurement method, which can be used for detecting various target substances, and can be widely applied in fields other than biomedicine such as environmental control and industry. . In order to achieve the above other objects, the present invention provides a mercapto-oligothiophene conductive derivable compound which is capable of detecting and measuring a power-measuring antibody probe. The electro-sensing sensor comprises a second electrode and a layer of antibody immobilized on the surface of at least one of the electrodes. Conductively-induced molecules are attached to and/or distributed between the electrodes to which the antibody is immobilized to break into electrical conductivity between the electrodes. In a particular aspect, the antibody probe molecule of the electrosensitized wafer of the present invention and related methods can be said to be "wearing a tight layer with electrical conductivity", which makes electrical conductivity in the system. Essentially becomes "amplified" to the extent that the instrument used today is sufficient for accurate sensing. The measurable electrical parameters of the electrical sensing wafer, such as resistance values, thus become not only detectable but also readable, which in turn becomes a meaningful parameter for interpretation. The present invention provides an electrosensor that achieves sensing of the above and other target analyte antigens. The electrosensitizer comprises two separate and uncoupled electrodes, and a layer of antibody immobilized on the surface of at least one of the electrodes, and is capable of binding to a highly specific antigen. When the mixed buffer of the sample to be tested flows through the surface of the wafer, the antibody contacts and binds the antigen, thereby changing the conductive characteristics of the original electrode, i.e., displaying a signal change difference represented by electrical sensing between the electrodes. Another specific evidence of the present invention is to provide an electrical sensor that achieves the sensing of the above and other target analytes 201200873. The electrical sensor includes two separate and unconnected electrodes, and is fixed to At least one layer of antibody on the surface of the electrode and can react with a highly specific antigen. The conductivity of the antibody electrode is promoted by the conductive attracting molecule, and the electrical conductivity between the two electrodes is improved. When the mixed buffer of the sample to be tested flows through the surface of the wafer, the antibody contacts and binds to the antigen, thereby changing the conductive characteristics of the original electrode, that is, displaying the difference in signal change represented by the electrical sensing between the electrodes. [Simple Description of the Drawings] Figure 1 shows the architecture of a basic electro-sensing system. 2A and 2B show two possible configurations of an electro-sensing wafer. Figures 3A and 3B illustrate how conventional electrophysiological measurements are not suitable for antigen detection using antibody probes. Figures 4A-4C respectively illustrate the preparation of a preferred embodiment of a sensing wafer in accordance with the present invention and the sensing of a sample thereof. Figures 5A-5C respectively illustrate the preparation of another preferred embodiment of a sensed wafer in accordance with the present invention and the sensing of a sample thereof. Figure 6 illustrates why the electro-sensing wafer and method of the present invention have substantial utility. Figures 7A and 7B show two examples of surface modification procedures for sensing wafers of the present invention which can enhance electrical conductivity throughout the detection system. Figures 8-10 show the results of several sample tests using the electro-sensing device of the present invention, respectively. Figure 11 illustrates the physical construction of an embodiment of the electro-sensing wafer of the present invention. Figure 12 shows a flow path configuration of an electro-sensing wafer handler in accordance with an embodiment of the present invention. [Embodiment] The present invention utilizes the electrical conductivity of the sensing wafer system (the wafer environment and the antibody molecules itself reacting with it) to achieve a practical use of the inductance measurement. In particular, the antibody probe molecule of the sensing wafer of the present invention and related methods can be said to be "wearing a tight-fitting electric conductive body", which makes the electrical conductivity in the system substantially become It is "enlarged to the extent that the instrument of today is sufficient for accurate sensing. Sensing the measurable electrical parameters of the wafer, such as resistance values, so that not only becomes detectable and can be interpreted, it becomes A meaningful parameter for interpretation. According to the present invention, an antibody immobilized on a sensing wafer as a detection probe is substantially converted into a semiconducting or even a good electrical conductivity by its non-conducting nature. Allowing the value of the electrical impedance change of the target substance 201200873 (after contact with the antibody on the sensing wafer and reacting) in the sample liquid to be tested, not only becomes detectable by the instrument, but can be measured with sufficient accuracy. Discrimination. The measured reading value can therefore be applied to the interpretation of the original target substance detection. In fact, as the art can be used by the skilled person The solution, in addition to the impedance value, other electrical parameters such as the capacitance value of the system, is also due to the improvement of the electrical conductivity of the entire system, and all of them can be measured. In addition, as an interpretation of the resistance value In addition to the rigorous definition of the reciprocal, the term "conductivity" is also more widely interpreted within the scope of the invention as the electrically conductive state in its system. Therefore, the term "conductivity induced" should be interpreted herein as "the improvement of the electrical conduction state. Thus, the sensor and method of the present invention can establish an electrically conductive environment that can be tolerated by being The captured target substance appears in the environment, and the resulting change in electrical conductivity becomes not only detectable to the IJ, but also can be measured and interpreted. Since the sensor and method of the present invention are substantially Amplifying the electrical characteristic detection range of the entire test sample system, so any change in electrical properties, whether impedance > current or capacitance, whether measured by direct current (DC) or any selected frequency of alternating current (AC) It can be easily detected and accurately quantified, and the amount of change can be a quantitative indicator of the amount of target substance present in the sample to be tested. Figure 1 shows a basic electro-sensing system of the present invention. The sensing wafer 100 constructed on a substrate 110 has an entire layer of antibody probes 12G fixed on the surfaces of its positive electrode 112 and negative electrode 114. The electrode therein may be, for example, Au. A metal thin film of Ag, Cu or Ni. The electrodes 112 and 114 serve as a substantial substrate for the antibody probe selected to be immobilized as a specific antigen. An embodiment of the novel electro-sensing technique of the present invention based on the electro-sensing wafer 100 The system, in combination with a detection circuit and a fluid system, provides a sensing chamber 102. Within the chamber, the test sample is brought into contact with the wafer to allow the target antigen molecule 134 suspended in the liquid sample to be It becomes the antigen 132 captured by the antibody probe 120. As will be described in detail below, the system of Figure 1 allows accurate measurement of the target antigen concentration in the test sample. This is utilized in the electrode of the sensing wafer. A voltage is applied between and measured by a current measuring instrument, as shown in the figure. Figures 2A and 2B show two possible configurations of an electro-sensing wafer according to a preferred embodiment of the present invention (configuration) The sensing wafer 200A of Fig. 2A is in the form of a typical flat wafer in which the sensing electrodes 212A and 214A are placed side by side on the substrate 210A. The flat sheet-shaped wafer structure, Relying on the cooperation of its corresponding wafer reading processing device to form a sample cavity in which sensing can be performed. 201200873 In contrast, the electro-sensing wafer 200B of FIG. 2B has a tubular configuration, The two sensing electrodes 212B and 214B are disposed at their mutually facing positions on the inner surface of the tubular "substrate j 210B. With such a tubular configuration, as long as the wafer is inserted into its corresponding processing and reading machine, the two are The end is closed, and the sample cavity 202B is easily obtained by sensing the wafer 200B. Figures 4A-4C respectively show the preparation of a preferred embodiment of the sensing wafer according to the present invention and the sensing of a sample thereof. . It is noted that the electrodes, antibodies, antigens, and conductivity-exposed molecules shown in the figure are not shown in the correct scale. In order to facilitate the explanation of the present invention, some of the figures shown in the drawings are shown in an exaggerated proportion. Fig. 4A shows the basic system of a sensing wafer of the present invention in which the overall conductive properties are enhanced by the use of conductive attracting molecules in an electrically conductive environment. In a preferred embodiment, a film-form Au is used to form substantially positive and negative electrodes 412 and 414 on substrate 410 of sensing wafer 400. Metals such as Ag, Cu, and Ni are also utilized in the present invention to fabricate electrodes. Suitable alloys, such as indium tin oxide (ITO), may also be used depending on the application. The electrically conductive molecule 442 is bonded to the electrode, as shown in the figure, which is attached to the electrode surface. According to the invention, such molecules thus form conductive attracting molecules immobilized on the surface of the electrode. When the wafer of the present invention is used, this allows the basic sensing system of the present invention to provide an enhanced electrical conduction environment, since the conductive attracting molecules modify the surface properties of the sensing wafer, As a result, the electrical conductivity of the bare sensing wafer system is improved. That is, after the antibody probe molecules are present in the wafer system, the conductivity between the positive electrode and the negative electrode is greatly enhanced. Because of the greatly improved electrical conduction environment between electrodes 412 and 414, as indicated by 405A in the figure, the system can sense current changes before and after the presence of antigen between the electrodes of sense wafer 400. Substances suitable for use as conductive attractants include, but are not limited to, oligothiophene-silane, oligothiophene-thiol, 1-phenyl oligothiophene ((l-phenyl)- Oligothiophene), 2-phenyl-oligothiophene, side-arm oligothiophene, oligophenyl oligothiophene, and derivatives thereof. In Fig. 4B, an antibody 422 for probe use of a specific target antigen is applied to the conductively induced molecular layer of the sensing wafer 400 and covalently bonded thereto. Because of this layer of immobilized antibody, at this stage (ie, when the target antigen has not yet appeared), the conductivity of the sensing wafer is somewhat reduced in this already electrically conductive environment 405B, but is still easy to use. The instrument is within the range of measurement. When antibody 422 is present, wafer 400 of Figure 4B forms an electric antibody sensor ready for electro-sensing applications for its particular target antigen. For any of the pre-set 201200873 sensing applications, the corresponding non-conductive antibody molecules need to be immobilized on the wafer first. For example, an immunoglobulin G molecule can be used as an antibody probe for detecting, for example, S100, alpha-fetoprotein, and tropolin I. The overall electrical conductivity of the system is reduced, to the extent that it reflects the presence of the probe in the system. This change in electrical conductivity becomes the reference value of the measurement at the time of detection. Fig. 4C shows the case where electrosensing is performed by exposing a target antigen to a probe antibody immobilized on a wafer. The wafer 400 in Figure 48 in which the probe has been prepared for sensing is exposed under a test sample. Since the probe antibody 422 that can be detected for a specific target has been immobilized on the wafer, the target existing in the sample, that is, the antigen 432, is captured by the antibody or, in other words, interacts with the antibody. (interaction) reaction. As the captured antigenic molecules 432 appear in the system, the overall conductivity of the entire electrically conductive environment 405C changes further (i.e., compared to Figure 4B), and its electrical impedance measurement reads. The change (i.e., measured as the current between the electrodes) becomes an indicator of the extent of the amount of antigen present in the system. According to the electro-sensing test performed by the present invention, when a sample containing a non-conductive target antigen is introduced into the fluid detection measurement environment provided by the sensing wafer of FIG. 4C, the overall electrical conductivity of the system is followed. The change is reduced. This reduction is reflected by the corresponding decrease in the measured current value. The degree of variation is proportional to the number of antigens that are captured by the wafer. However, it should be noted that in some cases, after binding of certain antigens in the test sample to the antibody probe, it may lead to conductance compared to the fact that the antigens have not appeared in the system before they appear in the system. The increase in degrees. Figures 5A-5C respectively illustrate another preferred embodiment of sensing a wafer in accordance with the present invention, which is prepared and sensed for a sample. The examples depicted in Figures 5A-5C and those shown in Figures 4A-4C, except for the physical configuration of the sensing wafers, are substantially identical except that they are positioned with electrodes facing each other. In accordance with the inference of the present invention (but the invention should not be limited to this inference), the arrangement of such electrodes may be due to their preferred conductive properties relative to the planar configuration of Figures 4A-4C. It is allowed to produce better sensing performance. Figure 6 is a schematic diagram showing the practical use of the electro-sensing wafer and method of the present invention. The graph in the graph shows the relative nature of the conductivity of a test sample compared to the concentration of antigen present in the sample. The vertical axis in Figure 6, the axis of electrical conduction properties, on the symbols A, B, C, D, D' and D", respectively, is the electrical conductivity of the sensing wafer at various stages of its preparation: A: Substrate B: Electrode C: Conductivity induced 201200873 D, D', D〃: Antibody probe addition In order to perform a wide range of concentration measurement on the target antigen present in the sample, the conventional technique is shown in the figure. The small current measurement reading range (BD^ or BD", depending on whether the probe to be added will slightly decrease or increase its overall electrical conductivity) attempts to measure the conductivity reading of the sample. The current value reading range is too small to be practical, and it is difficult to judge whether the target substance exists or not, and it is not necessary to obtain an antigen concentration curve in the sample with an acceptable reading resolution, E' or E". In contrast, if a conductive attracting molecule is used according to the present invention, the detection range (BD) of the target substance is substantially enlarged in the sense of a certain level, which can thus allow a good resolution. In other words, the accuracy is better, and the target concentration is read. This is because, as clearly shown by the characteristic curve E in Fig. 6, regardless of whether the concentration of the target substance in the liquid sample environment is linear or nonlinear between the corresponding measured currents, the broad measurement range is corresponding. The target detection carried out inside can of course make the interpretation of the instrument readings easier. 7A and 7B are two examples of surface modification procedures for sensing wafers of the present invention that can improve electrical conductivity throughout the detection system. Figures 8-10 show the experimental results of several sample tests using the sensor of the present invention, respectively. Figure 8 shows the binding curve of the S100 antigen in the experiment, relative to the anti-S100 monoclonal antibody, at a measurement voltage of -0.2V. A sensing wafer using an anti-S100 monoclonal antibody (mAb) as a probe was used to detect the S100 antigen. The binding curve in Figure 8 is obtained under a detection voltage of -0.2 V, which shows a clearly readable relationship and will target the target substance in a whole sample per milliliter (millilitei 0 to 200 micrograms). The concentration range corresponds to a sensing current scale of 0 to 50 nano-ampere. Bovine serum albumin (BSA) was used as a control group sample in this experiment. Figure 9 shows a type A fetus Alpha fetoprotein in the experiment, under the measurement voltage of -0.5V, relative to the anti-alpha fetoprotein monoclonal antibody, the binding curve at different antigen concentrations. An anti-alpha fetoprotein monoclonal antibody (mAb) is used as a probe for sensing wafers and is used to detect alpha-type fetal protein antigens. The binding curve in the figure is at -0.5V detection voltage. The sample is obtained at a concentration ranging from 0 to 1 microgram per milliliter. This experiment shows a clearly readable relationship curve that corresponds to the concentration range of the target substance in the sample. On the sensing current scale of 0 to 500 nanoamperes, BSA was used as a control group sample in this experiment. Figure 10 shows anti-(tropolin I) monoclonal antibody In the experiment, under the measurement voltage of -0.2 V, the binding curve relative to cardiac tropolin I (tropolin). Using anti-(tropolin I) monoclonal antibody (mAb) Sensing Crystal as a Probe 201200873 Tablets were used to detect cardiac tropolin I (tropolin) antigen. The binding curve in the figure is below the detection voltage of -0.2V, ranging from 〇 to 200 μg per ml. The test sample was obtained. This experiment shows a clearly readable relationship curve that corresponds to the target material concentration range in the sample to the sensing current scale of 0 to 30 nanoamperes. BSA is used in this experiment. As a control group sample, Fig. 11 shows a physical configuration of an embodiment of the sensing wafer of the present invention. Among the examples of the flat wafer structure, the sensing wafer 1100 having the planar electrode structure has a glass substrate Ilia plated thereon. With electrodes 1112 1114 and 111⁄2 And edge connection electrodes 116Z Π 64 and 1166» similar to those commonly found in printed circuit boards. In this example, the electrodes include an intermediate reference electrode 1116 between the positive 1112 and negative 1114 electrodes of functional properties. Some edge-connected electrodes can be reserved for future new features. A typical metallic conductive coating for sensing the functional electrode application of the wafer includes Au, Ag, Cu, Ni, etc., or it may be a soft material electrode or other electrically conductive material. Typical inert metal layers available for reference electrodes include Pt and others. The wafer can be electrically connected to the edge of a sensing circuit, and the connection electrode can be commonly used in PCB technology. These sheets typically have a thickness of about 2,000 angstroms. Figure 12 shows the overall system architecture of an electro-antibody sensor in accordance with an embodiment of the present invention. In this sensor example, two liquid flow paths 1272 and 1274 are used. When the wafer 1200 is sensed, it can also be selectively When it is carried on a wafer stage 1208, when it is applied to the detection of a specified target substance (antigen), the sample is loaded into the sample circuit 1282 by the injector 1284 via the loading path. The sample is pumped by a buffer stored in the reservoir 1286 to sense the surface of the wafer electrode by using a buffer 1127. The detection sample and the buffer are then discharged and then discharged. The waste liquid tank 1288 is to be processed later. It should be noted that, for the novelty of the present invention, the present invention is only described in detail with respect to the preferred embodiments described in the drawings. In the system, one of several possible ways of performing conductivity trapping, that is, the conductive attracting molecule is also used as a linker of the antibody counter electrode, that is, as depicted in Figures 4 and 5. Like Xi Ben It will be understood by the artist that other application arrangements of the conductive attracting molecules in the system of the invention are equally possible and feasible. For example, a slightly different arrangement includes the use of conductive deriving molecules to modify the antibody. Adding conductive attracting molecules to the buffer flowing through the surface of the sensing wafer, or adding it to the sample, or even any combination of the above arrangements, which are all feasible as confirmed by experiments. The present invention has been described in its entirety, and the various modifications, variations and equivalents thereof are still possible. The formula should not be construed as a limited interpretation of the scope of the patent scope of the invention.