201142287 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齊其相關方法。 [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)已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10, 1至100,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)是電學性的檢査。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 201142287 之變化量。 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，賣際上仍是極為不良，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，本發电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撿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。 4 201142287 本發明的另一項具體證據，此感測器系統中的抗體進一步與導電性誘昇分子連結。 再將抗體固定於至少一電極表面，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。 [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樣本進行感測 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樣本進行感 測之情形》 圖ό解釋本發明電性感測晶片及方法何以具有實質用途。 圖'10說明本發明感測晶片透過導電性誘昇分子懸浮於緩衝液中之實例，其可以提昇 整個檢測系統中的電傳導性。 圖11顯示本發明中抗體與導電性誘昇分子連結之架構。 [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」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(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 201142287 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」。 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到，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Π4表面上的整層的抗體探針120,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(chambery 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。 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」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。 201142287 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-砂焼(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，並與其共價鏈鍵結。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己具有電導性的環境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 201142287 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4Β中己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已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 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 201142287 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7-10示意圖顯示本發明感測晶片使用含有導電誘昇分子懸浮緩衝液進行表面修飾程 序，針對圖7所示之空白晶片，電性感測完全依賴誘昇分子742引入本系統，並將此抗 體分子722固定於電極712/714之晶片700表面，裸露於晶片表面並將待測物檢體之緩 衝液注入感測腔體702。 反之，對於分別在圖8, 9及10中之系統晶片800, 900和1000，實現以大致相同方式 進行電性感測，且依賴導電性誘昇分子帶入該系統之緩衝液中。圖8, 9及10中描繪該 系統之差異並與圖7比較。圖7之感測晶片是將導電性誘昇分子固定於晶片表面上。圖 8之感測晶片800是將導電性誘昇分子842與抗體分子鍵結822。而對於圖9之感測晶 片90Q抗體922必需固定在含有導電性誘昇分子之作為連結劑942之電極表面上使用。 可在待測物進行測試前，將含有導電性誘昇分子之緩衝液引入。至於圖10之感測晶片 1000,每個抗體分子1022均含有導電性誘昇分子1042鍵結在其上。 圖11說明更詳細的抗體1122與導電性誘昇分子1142共價鍵結之相關情形。抗體，類 似具有Υ形之分子機構定義為1122且含多重導電性誘昇分子1142鏈結其上。此類導電 性誘昇分子，如寡苯基-噻吩及其衍生物可由圖解說明，將噻吩分子11421修改其末端 官能基成苯基1142Ζ並共價鍵結在抗體上1122»部份導電性誘昇分子，定義為1142Α, 可在末端延長並結合抗體1122之Υ型分子於電極晶片上，且同時增進電導度。 上述是一描述完整的具體表現，可應用於包含各種表面修飾，可替代性架構及其平衡. 因此，本發明範圍不被上述之描述及圖說所限制。201142287 6. DISCLOSURE OF THE INVENTION [Technical 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 a sensor that can be used for the detection and measurement of electro-acceptable antibody probes. [Prior Art] In the field of medicine and related applications, biochips are used (biochip) It is a common practice to detect the target substrate in the test sample. Biochip sensors (or biosensor) have been applied under considerations such as accuracy and cost. Detection of a specific target substance. If the technology is feasible and the cost is acceptable, for any imaginable use, if the quantitative detection is performed in addition to the presence or absence of the detection of the target substance, the biochip sensing Obviously, it is more useful. For example, in biomedical applications, if, for example, from 1 to 10, 1 to 100, or even higher resolution, and Accuracy, while measuring the extent to which the target substance appears in the sample, is clearly informative for its intended sensing application. Biofilm based on optical sensing technology, It is common for today's global biosensing technology. This type of wafer relies on the need to use large, bulky and expensive precision optics for optical sensing to read the results of the test on the wafer. For these problems, bio-wafers using the principle of electro-pressure are obviously a reasonable method of miniaturization and low cost. In the technology of electro-sensing wafers, after the wafer is placed in the test sample, it is inspected (or sensed). It is an electrical inspection. The sensed on the sample to be tested may be impedance, capacitance, resistance, conductance, current, or any other. Any electrical parameters that are useful. However, until now, electrical biosensing techniques have limitations, due to the vast majority of liquid bioassays, None of them are electrically conductive. Figures 3A and 3B illustrate how conventional electro-sensing techniques are not suitable for antigen detection using antibody probes. For example, Figure 3A illustrates a prior art sense of the art. The test wafer 300 is, 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 may be, for example, gold (Au) or silver (Ag). ), a metal film of copper (Cu) or nickel (Ni). For practical use, the system must be capable of accepting an environment generally indicated by reference numeral 305 in the figure, i.e., sensing the amount of change between current electrodes 201142287 between the two electrodes of wafer 300. However, after the introduction of the sample to be tested, after the antigen molecule 332 therein (which is mostly non-conductor or 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 poorly sold, 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 Therefore, the purpose of the present power generation is to provide for the detection and measurement of electrical sensing antibody wafers for detecting 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-scale 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. 4 201142287 Another specific evidence of the invention is that the antibodies in the sensor system are further linked to a conductive eliciting molecule. The antibody is then immobilized on at least one electrode surface and can be reacted 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. [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 the same. Figures 5A-5C respectively illustrate the preparation of another preferred embodiment of a sensing wafer in accordance with the present invention and the manner in which it is sensed. The Figure illustrates the practical use of the electro-sensing wafer and method of the present invention. Figure 10 illustrates an example of the sensing wafer of the present invention suspended in a buffer through a conductive attracting molecule which enhances electrical conductivity throughout the detection system. Figure 11 shows the structure of the antibody in the present invention linked to a conductive attracting molecule. [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 The instrument that is "zoomed in" to the application of today is sufficient for accurate sensing. Sensing the quantifiable electrical parameters of the wafer, such as resistance values, so that not only becomes detectable and can be interpreted, it thus becomes a meaningful parameter for interpretation. According to the present invention, an antibody immobilized on a sensing wafer as a detecting probe is substantially converted into a semiconducting or even a good electrical conductivity by its non-conducting nature. This allows the value of the electrical impedance change of the target substance in the sample liquid to be detected (after it contacts and reacts with the antibody on the sensing wafer), which can be detected not only by the instrument but also with sufficient accuracy. Test discrimination. The measured readings can therefore be applied to the interpretation of the intended target substance. In fact, as can be understood by those skilled in the art, in addition to the impedance value, other electrical 201142287 parameters such as the capacitance value of the system are also improved by the present invention by improving the electrical conductivity of the entire system. It can be measured. Furthermore, in addition to the rigorous definition of the reciprocal of the resistance value, the term "conductivity" is more widely interpreted within the scope of the invention as the electrical conduction state in its system. Therefore, the term "conductivity induced" should be interpreted as "an increase in the electrical conduction state". Therefore, the sensor and method of the present invention can establish an environment of electrical conduction, which can allow the change of electrical conductivity caused by the appearance of the captured target substance in the environment, and can be detected not only Measured and can be measured and interpreted. Since the sensor and method of the present invention substantially amplifies the electrical characteristic detection range of the entire test sample system, any variation in electrical properties, whether impedance, current or capacitance, whether it is DC or DC ) or any selected frequency of alternating current (AC) measurements are easily detectable and accurately quantified. The amount of change can thus be a quantitative indicator of the amount of target material present in the sample being tested. Figure 1 shows the architecture of 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 120 fixed on the surfaces of its positive electrode 112 and negative electrode Π4, wherein the electrodes may be, for example, Au, Ag, Cu or Ni, etc. Metal film. Electrodes 112 and 114 serve as substantial substrates for antibody probes selected for immobilization to a particular antigen. An embodiment system of the novel electro-sensing technique of the present invention based on the electro-sensing wafer 100 can be combined with a detection circuit and a fluid system to provide a sensing cavity (chambery 102. In this cavity, detection The sample is brought into contact with the wafer to allow the target antigen molecule 134 suspended in the liquid sample to become the antigen 132 captured by the antibody probe 120. As will be explained in more detail below, the system of Figure 1 allows for the detection of the sample. The target antigen concentration is accurately measured. This is measured by applying a voltage between the electrodes of the sensing wafer and passing through a current measuring instrument, as shown in the figure. Figures 2A and 2B show the basis. Two possible configurations of the electro-sensing wafer of the preferred embodiment of the present invention. The sensing wafer 200A of FIG. 2A is in the form of a typical flat wafer. The sensing electrodes 212A and 214A are attached to the substrate 210A. Side-by-side placement. Such a flat sheet-shaped wafer structure relies on the cooperation of its corresponding wafer reading processing device to form a sample cavity in which sensing can be performed. The electrical sensing wafer 200B of 2B has a tubular configuration, and the two sensing electrodes 212B and 214B are disposed at mutually facing positions of the inner surface of the tubular "substrate" 210B. The configuration is as long as the wafer is inserted into its corresponding processing and reading machine to close both ends thereof, and the sensing wafer 200B can easily obtain a sample cavity 202B. 201142287 Figures 4A-4C respectively show sensing in accordance with the present invention. A preferred embodiment of the wafer is prepared and sensed for a sample. It is noted that the electrodes, antibodies, antigens, and conductive attracting molecules, etc. shown in the figures are not drawn to scale on the correct scale. The explanation of the present invention is partially shown in an exaggerated scale. Figure 4A shows the basic conductive system of a sensing wafer of the present invention, in an electrically conductive environment, the overall conductive properties, This is enhanced by the use of conductively induced molecules. In a preferred embodiment, the film-form Au is used to form substantially positive and negative electrodes 412 and 414 on the substrate 410 of the sensing wafer 400. Metals such as Ag, Cu, and Ni can also be used in the present invention to produce electrodes. Depending on the application, a suitable alloy such as indium tin oxide (ITO) can also be used. Conductive molecules 442 are bonded to the electrodes, as shown in the figure, which are immobilized on the surface of the electrode. According to the present invention, such molecules thus form a conductive attraction immobilized on the surface of the electrode. Molecules. 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 because the conductive attracting molecules modify the surface 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 the change in current before and after the presence of the 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 Figure 4B, antibody 422 for probe use of a particular antigen of interest is applied to the layer of conductively induced molecules of the sensing wafer 400 and is covalently linked 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 the already conductive environment 405B, but it 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 molecule. For any predetermined sensing application, the corresponding non-conductive antibody molecule corresponding to it needs to be first immobilized on the wafer. 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 reference for the measurement at the time of the 201142287 test. Fig. 4C shows a case where electrosensing is performed by exposing a target antigen to a probe antibody immobilized on a wafer. The wafer 400 in Figure 4 that has prepared probes 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 present 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: Conductance induced D, D', D": Antibody probe addition In order to perform a wide range of concentration measurements on the target antigen present in the sample, the known technique is shown in the figure. The small current measurement reading range (BD' or BD 〃, depending on whether the probe being 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. 7-10 are schematic views showing the surface modification process of the sensing wafer of the present invention using a conductive excitation molecular suspension buffer. For the blank wafer shown in FIG. 7, the electro-sensing is completely dependent on the attracting molecule 742 to be introduced into the system. The antibody molecule 722 is immobilized on the surface of the wafer 700 of the electrode 712/714, exposed to the surface of the wafer, and the buffer of the analyte to be tested is injected into the sensing cavity 702. Conversely, for system wafers 800, 900, and 1000, respectively, in Figures 8, 9 and 10, electrical sensing is performed in substantially the same manner and dependent on conductive trapping molecules into the buffer of the system. The differences in the system are depicted in Figures 8, 9 and 10 and compared to Figure 7. The sensing wafer of Figure 7 is to immobilize conductive attracting molecules on the surface of the wafer. The sensing wafer 800 of FIG. 8 is a bond 822 of the conductive deriving molecule 842 to the antibody molecule. For the sensing wafer 90Q antibody 922 of Fig. 9, it must be immobilized on the surface of the electrode as the bonding agent 942 containing the conductive attracting molecules. A buffer containing a conductive derivatizing molecule can be introduced before the test object is tested. With respect to the sensing wafer 1000 of Figure 10, each antibody molecule 1022 contains a conductive deriving molecule 1042 bonded thereto. Figure 11 illustrates a more detailed description of the covalent bonding of antibody 1122 to conductively-exposed molecule 1142. The antibody, a molecular structure similar to a scorpion, is defined as 1122 and has a multi-conductive attracting molecule 1142 attached thereto. Such conductively induced molecules, such as oligophenyl-thiophene and its derivatives, can be illustrated by modifying the thiophene molecule 11421 to its terminal functional group to phenyl 1142 and covalently bonded to the antibody. The molecule, defined as 1142 Α, can extend at the end and bind to the Υ-type molecule of antibody 1122 on the electrode wafer while enhancing electrical conductivity. The above description is a complete description of the specific embodiments and can be applied to various surface modifications, alternative structures and their balances. Therefore, the scope of the present invention is not limited by the description and illustration.