Field of the invention The present invention relates to development of integrated enzyme-metal nanoparticle nano-architecture based biosensors for detection of glucose, uric acid and cholesterol at physiological level. The present invention also relates to sensing of hydrogen peroxide, uric acid, cholesterol and glucose using nanoarchitectured sensor. Background and prior art of the Invention Development of a reliable method for the sensing of H202 is of great interest in many different areas. H202 is the most valuable marker for oxidative stress; oxidative damages resulting from the cellular imbalance of H202 and other reactive oxygen species generated from H202 are connected to aging and severe human diseases like cancer and cardiovascular disorder. H202 is recognized as one of the strongest oxidant in the conversion of dissolved S02 to H2S04 is recognized as one of the strongest oxidant in the conversion of dissolved S02 to H2S04, which is the main contributor to the acidification of aqueous phase of the troposphere. More importantly, H202 is a product of the reaction catalyzed by a large number of oxidase enzymes. The operating principle of the oxidasebased electro-chemical biosensor involves the determination of enzymatically generated H202. Highly sensitive and selective transducer is very essential for the precise monitoring of H202 generated during the enzymatic reaction. Such sensitive transducer can, in principle, used in the development of clinically important analytes such as uric acid, cholesterol and glucose using the corresponding oxidase enzymes. The carbon paste, carbon nanotube, polycrystalline Pt. Mediator and peroxidase based electrodes have been used for the sensing of H202 (Anal. Chern. 1992,64, 1285; Langmuir 2005,21,3653; Anal. Chern. 2004, 76, 474). Deactivation of electrode surface due to the formation of surface oxide is one of the problems with the polycrystalline Pt electrode (anal. Chem.2003, 75, 2080). Many attempts have been made to improve the performance of Pt based electrodes. The Pt-black, Pt-black rnicroarray and mesoporous Pt electrodes have been used to achieve high sensitivity towards H202 (Talanta, 2006, 68, 1632, Anal. Chem.2002, 74,5717). The use of 111 icroelectrode array or nanoelectrode ensembles for electroanalytical applications has several advantages over the conventional rnacroeletrodes. In principle, the electroanalytical detection limits at micro/nanoelectrode ensembles can be much lower than that of an analogous macrosized electrode due to the enhancement in the signal to noise (SIN) ratio (AnaI.Chem.1987, 59,2625). Quantification of cholesterol uric acid and glucose in clinical sample is very essential as the elevated concentration leads to life threatening diseases. The normal total cholesterol level in serum is 3 to 6.5 mM (Anal. Chem.1993, 65, 3258). Most of the cholesterol (2/3) in serum exists in the form of ester. The normal uric acid level in serum range from 240 to 520 ~lM and in urinary excretion 1.4 to 404 mM (J. Assoc. Off. Anal. Chem. 1987, 70, I). The physiological level of glucose is 3-8 mM (The New England J. Medicine 2005, 353, 1454). First and second generation biosensors have been proposed for the precise estimation of these analytes. In both the cases, oxidase enzymes are used. In the first generation biosensors, the estimation of enzymatically generated hydrogen peroxide is considered to be the direct measure of the actual concentration of the analyte. On the other hand, oxidase enzymes along with a suitable mediator, which shuttles the electron transfer between the enzymes and electrode surface, is used in the second generation biosensors. The major problem associated with the mediator-based biosensors is the lack of long term stability due to the leaching of mediator from the electrode surface. Various methodologies have been used for the electrochemical sensing of H202 uric acid and cholesterol. I-b02 can be conveniently quantified electrochemically by its reduction or oxidation. Enzymes and redox mediators that catalyze the oxidation/reduction I-b02 have been conventionally used. Peroxidases, Prussian Blue, carbon nanotube (CNT), mesoporous Pt microelectrodes and metal particle-based electrodes are widely used (AnaI.Chem.2004, 76,474; Anal. Chem.2002, 74, 1322; Biosensor Bioelectron. 2009, 24,3264). Detection of H202 by its reduction is susceptible to interference from oxygen. On the other hand oxidation of at more positive potential invites interference from other small electroactive molecules. One of the serious problems in the electrochemical detection of H202 is the interference due to easily oxidizable analytes commonly available in the samples. Mitigation of interference without compromising the sensitivity is a challenging task. The electrode that can catalyze the oxidation of H202 at less positive potential (DAV) is desired for the sensing of H202. The redox mediator-bases electrodes are not stable. presumably due to the leaching of the mediator from the electrode surface. Lyon and Stevenson achieved the lowest detection limit of 8 pM H202 using chemically ,.' activated redox mediator (Anal. Chem. 2006, 78, 8518). However, this method requires diffusional redox mediator and enzyme horseradish peroxidase. The disadvantage of this method is that the redox mediator easily absorbs on the electrode surface and hence deactivation of the electrode is inevitable. A detailed survey of literature on the electrochemical sensors for H202 is given in Table I. Jyh-Myng Zen and Jen-Sen Tang (Anal. Chem. 1995, 67, 1892) modi fled the glassy carbon electrode by Nafion !RU2_'Pb,07. x (ruthenium oxide pyrochlore) and detected uric acid by Osteryoung square -wave voltammetry. The linear range obtained with this· electrode was from 7.5 x 10.5 to 5x 10-7 M and the limit of detection was 1.1x 10-7M The major concern with this electrode is that it needs to be operated in acidic pH (pH I). Cai et al. utilized electrochemically activated carbon paste electrode for the detection of uric acid (Talanta, 1994, 41,407). This electrode showed linear response from 3x 10-8 to 2Ax I0-4 M and the limit of detection was 1.2x 10-8 M. However, the preparation of the electrode is rather time consuming and the linear range was not in the diagnostic range of the physiological sample. Yu et al. modified glassy carbon electrode with polyglycine and examined the voltammetric response towards uric acid and ascorbic acid (0.30 mY vs. SCE) and increased the response current of uric acid after redox reaction. Although the electrode could distinguish the voltammetric response, the linear range was out of normal physiological level. Nakaminami et al. developed enzyme based uric acid sensor using redox polymer and enzyme uricase. This method is based on the mediated oxidation of uric acid by the enzyme. The analytical parameters such as the limit of detection, linear range etc. was not reported (Anal. Chem. 1999, 71, 1928. Raj and Ohsaka utilized selfassembled monolayer of heterocyclic thiol for the non-enzymatice detection of uric acid (J. Electroanal. Chem. 2003, 540, 69). This electrode could sense uric acid in the presence of ascorbic acid. However, the lowest concentration detected was 1 x J 0.6 M. Uric acid and cholesterol are traditionally detected by time consumtng colorimetric methods. These methods require color indicating reagents and enzymes. Uric acid has been quantified in real samples using chemically modified electrodes and redox mediator based electrodes. Carbon-based electrode with preconcentration procedures has been widcly used for the voltammetric sensing of uric acid. Sol-gel derived ceramic film with redox mediator has been used for the sensing of uric acid by the oxidation of uric acid lAnaI. Chem.2002, 74, 5734-5741]. Redox mediator-based biosensors have been llsed for the sensing of uric acid (Anal. Chern. 1999,71,1928-1934; Anal. Chern. 1999,71,4278- 4283). ruj ishima and co-workers have reported the voltammetric sensing of uric acid using boron doped diamond electrode; this electrode requires as large as 0.9 V (SCE). Nanostructured carbon fiber electrode has been used for the sensitive detection of uric acid fast sweep voltammetry (AnaI.Chem.2000, 72, 1576-1584). Selectivity and sensitivity is a major problem with unmodified electrodes. The enzyme based electrodes highly selective and sensitive in the determination of uric acid. Unlike uric acid, direct oxidation of cholesterol is not possible on conventional electrodes. Traditionally, sensing of cholesterol has been performed with enzymes cholesterol oxidase and cholesterol esterase using redox mediators (Biosensors and Bioelectronics 2008, 23, 1083-1100 and the references cited therein). Lack of long term stability, selectivity and poor sensitivity are the draw backs with the existing sensors reported in the literature. Lu et at. (200 I - 818000) disclosed the development of oxidase base sensor for the sensing of uric acid using oxidase enzyme and metal doped carbon composite electrode. Hydrogen peroxide is generated during the enzymatic reaction of uric acid with enzyme. Enzymatically gen'erated hydrogen peroxide was electrochemically detected by its reduction. Shen et al. (1999 - 295400) developed non-enzymatic electrochemical sensor for the sensing of uric acid using water soluble redox mediator and polymer. Stability, lowest detection Iim it are not given. Hsiung et at, (2005-211679) disclosed the fabrication of enzyme based sensor for the sensing of uric acid using enzymes catalase, uricase and redox mediator ferrocene carboxylic acid. The amperometric current due to the reduction of enzymatically generated hydrogen peroxide was used to quantify the concentration of uric acid in the sample. Polyacrylamide was used to bind the enzyme and redox mediator. Chuang et al (2007 - 782517) disclosed a biosensor, a biostrip and a manufacture method of determination of uric acid by a non-enzymatic reagent. The biosensor of the invention comprises at least a biostrip and a non-enzymatic reagent for the determ ination of uric acid. The biostrip has at least two working electrodes and at least one reaction zone. The non-enzymatic reagent is disposed in the reaction zone' on the biostrip where a tetrazolium salt and an active electron mediator are immobilized. The reaction zone is used for a sample to initiate the oxidation-reduction rcaction of the tctrazolium salt and then, according to the oxidation - reduction reaction, with the active electron mediator and the two working electrodes an electronic signal is transmitted to the sensor to produce a corresponding micro-current intensity which is in turn calculated by thc sensor to reflect the uric acid level in the sample. Kumar et al. disclosed (US 7175746) the development of polymer based cholesterol biosensor using potassium ferricyanide as a mediator. The enzyme cholesterol oxidase was physically adsorbed on the electrode surface and it has a response time of 30 s. The real sample analysis has not been performed with the proposed sensor. The major concern with this biosensor is that it may not be used directly for the quantification of total concentration of cholesterol in real sample analysis, as it uses only cholesterol oxidase. In serum, 2/3 of cholesterol exists as ester and quantification of total cholesterol in serum requires the enzyme cholesterol esterase. US 7257837 disclose the development of cholesterol sensor using cholesterol oxidase, solgel silicate in the presence and absence of redox mediator. The enzyme and redox mediators were encapsulated in to the silicate net work and cast on the conducting substrate. In the absence of redox mediator, cholesterol was detected at the potential of 0.9 V by measuring the current for the oxidation of enzymatically generated hydrogen peroxide. Interference due to other small molecules was noticed, due to the high overpotential. In the prescnce of redox mediator, the detection potential was decreased to 0.4 V. This sensor has the response time of 30-90 s and was reused at least for five times. The utilization of the sensor for real sample analysis has not been demonstrated and the sensor may not be used for the quantification of total cholesterol in real samples. US 6117289 describe a cholesterol sensors fabricated using cholesterol oxidase, cholesterol esterase. redox mediator and a surfactant. The sensor operates at the potential of 0.5 V. The reaction layers contain the enzymes and the redox mediator. The response time of the sensor is 9 min and the utilization of surfactant may interfere in the real sample analysis. Real sample analysis has not been described. US 6071392 disclose a cholesterol sensor which comprises an electrode system having a working electrode and a counter electrode formed on an electrically insulating base plate. The electrode coating layer comprising water-soluble cellulose derivatives and saccharides for covering the electrode. The reagent layer contains the enzymes cholesterol oxidase and cholesterol esterase and surfactant. The configuration of the sensor eliminates impairment of sensor response due to electrode degeneration caused by invading surfactant into the electrode system. The sensor was operated at the potential of 0.5 V. US 62146] 2 disclose a method for the quantitative determination of cholesterol using a carbon working and counter electrodes. The reaction reagent contains cholesterol dehydrogenase, electron transfer mediator, nicotinamide adenine dinucleotide (NAD) and diaphorase. NAD is reduced to NADH during the enzymatic reaction; NADH oxidized to NAD by the electron transfer mediator. The concentration of the reduced form of the electron transfer mediator is directly proportional to the concentration of cholesterol. The reaction ofNADH with mediator was catalyzed by another enzyme diaphorase. All these reactions are favorable only at pH>9. The reaction is very slow at lower pH. The operating potential of the sensor was 0.5 V. For real sample analysis, high pH is not desirable. Apart from the above prior arts, others also discloses the sensing of cholesterol using one or more redox mediators, surface active reagents (US patents 6966977, 1239048, 5695947, 6451372,6821410,5695947,1347292,2007-0158189, 6117289, 2002-084937,1318396, 6117289,7267837, 1239048,6342364). However, the standing difficulties are: (i) require large overpotential, (ii) lack of selectivity, (iii) suitability of the sensors for real sample analysis etc. US 2004/0077844 (US'844 in short) relates to fabrication and arrangement of nanoparticles into one-dimensional linear chains achieved by successive chemical reactions, each reaction adding one or more nanoparticles by building onto exposed, unprotected linker functionalities. Nanoparticle spheres are functionalized in a controlled manner in order to enable covalent linkages. Functionalization of nanoparticles is accomplished by either ligand exchange or chemical modification of the terminal functional groups of the capping ligand. The method tor assembling nanoparticles in a controlled fashion comprises the steps of: providing a plurality of nanopariicles; providing a plurality of ligands. each of the ligands comprising at least one linker arm; attaching at least one Iigand to each of the plurality of nanoparticles; and reacting at least one pair of the linker arms to ronn an assembly of nanoparticles. Fadime et al (Sensors 2009) relates to a amperometric cholesterol biosensor with immobilization of cholesterol oxidase on electrochemically polymerized polypyrrole polyvinylsulphonate (PPy-PYS) films has been accomplished via the entrapment technique on the surface of a platinum electrode. Fadime et al. further reports that the immobil ization of cholesterol oxidase onto PPy-PYS film via an entrapment procedure for determination of free cholesterol. Fadime et al. used enzyme immobilized polymer based working electrode, Pt counter and Ag/ AgCI reference electrodes. Immobilization of cholesterol oxidase was carried out by the physical entrapment approach. The concentrations of pyrrole and polyvinylsulphonate were 0.1 M, respectively. The monomers were electrochemically polymerized on the electrode surface in the presence of cholesterol oxidase. Amperometric response studies were also carried out in phosphate buffer. Operational stability, storage stability, pH and temperature were determined via application of +0.4 Y with respect to Ag/AgCI electrode to detect oxidation current of H202• The study shows that uric acid has no interfering effects on the analysis of cholesterol. But interfering effects of ascorbic acid, paracetamo] and glucose on the analysis of cholesterol were observed. These interferences were almost removed by dilution of solution. Further, these abovementioned literatures describe the electrochemical sensing of cholesterol by enzymatic methods. Most of them involve the use of mediators. The enzymatically generated hydrogen peroxide was detected either by its oxidation or by its reduction. Oxidation of hydrogen peroxide at more positive potential invites interference form other analytes co-exist in the sample. Oxidation at less positive (<<;0.45V) is preferred without mediator. The second approach is by the reduction at negative potential: this approach is susceptible for interference due to dissolved oxygen. In many of the literature, the real sample analysis has not been performed to validate the utility of the sensor for the analysis with real systems. Stability, sensitivity and response time of the sensors need significant improvement. The sensor described in the aforementioned prior art does not deal with the real sample analysis and cholesterol ester. It describes the use of conventional polymer and cholesterol oxidase for the sensing of commercially available cholesterol (not total cholesterol in serum) which may not be applicable for real samples. Furthermore, ascorbic acid, paracetamol and glucose severely interferes the measurement. It is mentioned in Fadime et al. that dilution of the sample can remove the interference. Dilution of real sample is not recommended. Hence there is need for improved method for use in real samples that do not suffer from interference with other analytes and other drawbacks of the prior art and yet provide improved stability, sensitivity and fast response time of the sensors. Object of the invention The principal object of the present invention is to provide a biosensor by integrating Pt nanoparticles and oxidase based enzymes with a conducting substrate. Another object of the present invention is to provide an integrated enzyme-metal nanoparticle nano-architecture for the sensing of uric acid, cholesterol and glucose at physiological level. A further object of the present invention is to provide a fabricated biosensor that detects hydrogen peroxide at sub nanomolar level without any interference. Yet another object of the present invention is to provide a fabricated biosensor that detects uric acid/cholesterol/glucose without any redox mediator at physiological level. Yet another object of the present invention is to provide a sensor that does not require sophisticated equipments. Yet another object of the present invention is to provide a sensor that docs not suffer from interference from other electroactive interferents present in the real samples. Summary of the invention According to one aspect of the present invention. there is provided an integrated enzymemetal nanoparticle based sensor comprising integrated enzymes and nanoscale platinum particles h<:iving size ranging from 7 to 10 nm on a conducting support modified with biocomposite layer. According to another aspect of the present invention. there is provided a fabricated sensor prepared by the method comprising the steps of I. modifying a conducting support comprising polycrystalline gold electrode with a layer of biocomposite derived from 3-(mercaptopropyl)trimethoxy silane (MPTS); II. encapsulating the oxidase enzymes into the network; iii. self assembling the platinum nanoparticles on -SH groups of silicate network by chemisorption. Brief Description of the accompanying Figures Figure 1 illustrates the fabrication ofnano-architectured biosensor. Figure 2 displays the amperometric trace obtained for the oxidation of H202 on the Pt nanoparticulate based architecture. Figure 3 illustrates the interference free amperometric sensing of H202• Ascorbic acid, uric acid and paracetamol (0.1 mM each) were injected at regular interval. Figure 4(A) shows the amperometric response obtained for the biosensing of uric acid with the nano-architectured biosensor. Figure 4(8) shows corresponding calibration plots. Figure 5(A) shows the amperometric response obtained for the biosensing of cholesterol cster at the nano-architectured biosensors. Figure 5(8) is the corresponding calibration plots. Figure 6(A) illustrates the amperometric response obtained for the biosensing of glucose with the nanoarchitectured biosensors. Figure 6(8) shows the corresponding calibration plot Detailed Description of the Invention The present invention relates to the development of integrated nano-architectured sensors. In the present invention. metal based nanoparticulates have been self assembled and encapsulated into a network. According to one embodiment of the present invention, there is provided an amperometric sensor for the sensing of clinically important analytes such as hydrogen peroxide, uric acid, cholesterol and glucose is developed using nanosized Pt particles and oxidase based enzymes. A biosensor is developed by integrating Pt nanoparticles and oxidase based enzymes with a conducting substrate. The enzymatic reaction of clinically important analytes with the oxidase enzymes generates hydrogen peroxide. The amount of enzymatically generated hydrogen peroxide is directly proportional to the actual concentration of the analyte present in the sample. Enzymatically generated hydrogen peroxide is quantified amperometrically at the potential of 0.45 V (Ag/AgCl) by the Pt nanoparticles on the sensor. In the case of uric acid biosensor, enzyme uricases is used whereas enzymes cholesterol oxidase and cholesterol esterase are used in the development of cholesterol biosensor; glucose oxidase is used in the case of glucose biosensor. Real samples have been tested to val idate the method. According to the present invention, there is provided a sensor comprising a three dimensional (3D) silicate network on a solid support used for the integration of enzymes and Pt nanoparticle. Enzyme and Pt nanoparticle integrated 3-D network on a solid support has been developed for the sensing of total cholesterol and uric acid. The enzymes are encapsulated into the 3-D network and nanoparticles are chemisorbed onto the thiol groups. Such integration of Pt particle and enzyme(s) (uricase, glucose oxidase, cholesterol oxidase, cholesterol esterase) for the fabrication of electrochemical sensors has not been disclosed in prior art. The distribution of metal nanoparticle throughout the network is an added advantage; it favors the facilitated reaction. Different metals that can be used as conducting support are polycrystall ine Au. Pt, Pd and Cu. The present invention teaches a sensor having both cholestcrol oxidase and cholesterol esterase. It is well known that more than 80% of cholesterol in real sample (serum) exists in the form of ester, which cannot be detected only by cholesterol oxidase. Cholesterol ester should be hydrolyzed to cholesterol by another enzyme cholcsterol esterase. The hydrolyzed product (cholesterol) can undergo enzymatic reaction with cholesterol oxidase and generates hydrogen peroxide, which can be detected electrochemically. For the measurement of total cholesterol in serum sample, the sensor should have both cholesterol oxidase and cholesterol esterase. The present invention teaches the integration of enzyme(s) (cholesterol oxidase and cholesterol esterase in the case of cholesterol sensor) and nanoscale Pt particles on a conducting support. Such integration of Pt nanoparticles and enzymes of any kind has not been found either in the literature or in the prior art. The analytical application of the sensor is demonstrated with real sample analysis. The results are authenticated/validated with clinical laboratory measurements (as given in the Table 2). The sensor developed in the present invention could detect (i) hydrogen peroxide at sub-nanomolar level and (ii) cholesterol, uric acid and glucose at well below the physiological leve!' It has wide linear range and does not suffer from the interference due to other analytes. Dilution ofthc sample is not required in the present invention and more importantly it has the response time of 2s. further aspect of the present invention lies in the development of an integrated enzymemetal nanoparticlc nano-architecture for the sensing of hydrogen peroxide at subnanomolar level and sensing of uric acid, cholesterol and glucose at physiologicalleve!. Hydrogen peroxide is detected in range from 0.1 nM to 1.4 mM. Uric acid is detected 100 nM - 200 ~M, cholesterol is detected at 0.5 /-lM - 12 mM and glucose is detected at a range of 10 nM - 20 mM. The detection range can be varied by changing the enzyme loading in the sensor. The sensor can d~tect nanomolar levels (well below the physiological level) of glucose and uric acid and sub-micromolar level in the case of cholesterol. Further aspect of the invention involves the fabrication and characterization of the sensor/biosensor by using a conducting support modified with a biocomposite layer to form a 3D silicate network. Yet another aspect of the invention involves the formation of a self assembled group of the silicate network by chemisorptions comprising platinum nanoparticulates. Yet another aspect of the present invention features the experimental testing with commercial and real samples. The present invention provides a fabricated nano-architectured sensor which is stable for II days when stored in phosphate buffer solution of pH 7.2 having only 9% decrease in the initial response after 15 days. The operation stability of the samc electrode when subjected to 20 repetitive measurements for a period of 24 hI', the coefficient of variation in the current is only 0.12%. Advantages of the present invention The present invention unlike the existing system of the prior art has the following advantages. I. The present sensor is advantageous in view of sensing of hydrogen peroxide at sub-nanomolar level without any interference. 2. Biosensing of uric acid/cholesterol/glucose without any redox mediator at well below the physiological level and show linear response for wide concentration range. 3. The present invention does not require sophisticated equipments. 4. The present invention also does not suffer from interference from other electroactive interferents present in the real samples. The invention is now defined by way of non-limiting illustrative examples: Example I Fabrication and characterization of sensor/biosensor The fabrication of this sensor involves the following procedures: (i) the conducting support (polycrystalline Au electrode) was first modified with a layer of biocomposite (enzyme-sol-gel 3-D silicate network) derived from 3-(mercaptopropyl)trimethoxy silane (MPTS). This network has plenty of -SH functional groups. The oxidase enzymes were encapsulated into the network; (ii) Pt nanoparticulates have been self-assembled on the - SH groups of the silicate network by chemisorption. The size and morphology of the nanoparticulates on the silicate network have been examined with field emission scanning electron microscope (FESEM) measurement. The nanoparticulates are randomly distributed throughout the silicate network on the electrode surface and have the size distribution between 7 and 10 nm. FESEM images confirm the encapsulation of enzyme inside the silicate network. FESEM measurement was performed with lEOL lEM 6700F microscope. The Au coated glass slides were first functionalized with MPTS network and then Pt nanoparticles were immobilized by chemisorption (as depicted in Figure I). Example 2 The electrochemical measurements were performed with computer controlled CH 1643 electrochemical analyzer. Electrochemical cell consists of three electrodes. Working: nano-architectured electrode; auxiliary: Pt wire; reference Ag/ AgCI saturated with 3 M NaCI. The colloidal Pt nanoparticulates were synthesized according to the reported procedure (Chem. Commun. 2005, 2972 - 22974) with little modification. Briefly, 25 ml aqueous solution of H2PtCI6 (0.02 mM) and I mM glucose was mixed and stirred for 2 min. Then the pH of the solution was adjusted to 8 with NaOH. A 400 III of aqueous NaBH4 (0.05 M) was then added drop wise to the stirred solution: the stirring was continued for 15 min. The formation of nPt was followed by UV-visible spectral measurement. The color of the solution changed to grey-brown and the peak at 200-300 nm disappeared upon the addition of reducing agent, indicating the reduction of the PtCi6 2 - ions to metallic Pt. The MPTS sol was prepared by the dissolving MPTS, methanol and water (as 0.1 M HC!) in the molar ratio of I:3:3 and stirring the mixture vigorously for 30 minutes. For the fabrication of biosensor, the MPTS sol-enzyme biocomposite was prepared by the following procedure: first the MPTS sol was prepared by taking 24flL Hel (0.] M in 2.ml water and stirring the mixture vigorously for 30 min. Then 0.5 ml MPTS sol was mixed with 0.5 ml of enzyme (uricase) solution (20 mg/ml in 5 mM phosphate buffer solution of pH 8) and stirred for 2-3 min for the encapsulation of enzymes into the silicate network. For the fabrication of cholesterol biosensor, cholesterol, esterase (10 mg/ml) and cholesterol oxidase (20 mg/m!) were mixed with MPTS sol as described earlier. The resulting sol-gel biocomposite was stored at 4°C. The cleaned polycrystalline Au electrode was first soaked in MPTS sol-enzyme biocomposite for 20 min for the spontaneous adsorption of enzyme encapsulated sol-gel net-work on the electrode surface. The nPts were then self-assembled on the free thiol groups of the enzyme encapsulated sol-gel network at 4°C by soaking the enzyme encapsulated silicate network modified electrode (Figure I). Example 3 Amperometric sensing ofH202 Figure 2 displays the amperometric trace obtained for the oxidation of H202 on the Pt nanoparticulate based architecture. The electrode was polarized at 0.45 V and aliquots of H202 were injected into a stirred phosphate buffer solution (PBS). Fast and stable response was obtained upon every injections; the response time was 2s. The sensitivity and the experimental limit of detection (SIN = 9) were found to be 1.82±0.O I nA/nM and 0.1 nM respectively. The 3D nanoarchitecture design on the electrode surface allows the facile access of the reactant to the catalyst site and the rapid removal of the product by diffusion to solution from the film. The elimination of interference due to easily oxidizable compounds present in the physiological system is a real challenge in the voltammetric detection of H202 Ascorbic acid; uric acid and paracetamol are the major interfering agents present in the physiological system. To examine the performance of nano-architectured electrode in the presence orthe interferents, amperometric response of the electrode toward H202 has been tested in the presence of ascorbic acid, uric acid and paracetamol. The amperometric responsc of the nPt electrode was first registered by injecting H202 (30 M) into a stirred supporting electrolyte and an aliquots of ascorbic acid, uric acid and paracetamol (0.1 mM each) have been injected subsequently at regular interval (Figure 3). Interestingly, no observable change in the steady state current for H202 was noticed upon the addition of these interferents, indicating that the present electrode can be successfully used for the measurement of H202 without any interference. The sensor is highly stable, no change in the initial response was obtained for II days and 7% decrease in the current was noticed after 15 days (same electrode). Example 4 Measurement of H202 in rainwater The analytical utility of the Pt nano-architectured electrode was demonstrated by measuring the concentration of H202 in rainwater. Four samples were collected and analyzed amperometrically. Stable amperometric response was obtained upon every injection. Addition of standard H202 solution (standard solution means calibration plot was made with known concentration) showed linear relationship between the amperometric current and concentration of H202. The total concentration of H202 in the rainwater was in the range of 150 to 170 nM (Table 1). The recoveries for the spiked H202 (spiked solution means after the addition of rainwater, known concentration of H202 was spiked and the recoveries were calculated from the response) sample wcre 92-94%. Table 1: Determination of H202 in rainwater using nano-architectured platform a Sampleb 2 Original value Total concentration (nM) (nM) 1.5 150 1.66 170 1.68 186 1.66 166 3 4 a three independent amperometric measurements were made with each samples. b collected in the months of August 2007 (samples I and 2) and January 2008 (samples 3 and 4). Example 5 Biosensing of uric acid, cholesterol and glucose. The oxidase enzymes catalyze the oxidation of variety of analytes in the presence of oxygen; H202 is generated during the enzymatic reaction. Since the amount of enzymatically generated H202 is proportional to the analyte concentration, the H202 sensitive electrodes can be used for the measurement of analyte present in the biosample. The biosensor was fabricated by integrating the oxidase enzyme and nPts with the silicate network (Figure I) . .In the case of cholesterol, cholesterol esterase and cholesterol oxidase were used; cholesterol esterase hydrolyzes cholesterol ester into cholesterol and cholesterol oxidase converts the enzymatically generated cholesterol to 4-cholesten-3-one. The enzyme encapsulated into the network efficiently catalyzes the oxidation of uric acid / cholesterol! glucose in the presence of oxygen. Enzymatically generated H202 sensitive electrodes can be used for the measurement of analyte present in the biosample. The biosensor was fabricated by integrating the oxidase enzyme and nPts with the silicate network (Figure 1). In the case of cholesterol, cholesterol esterase and cholesterol oxidase were used; cholesterol esterase hydrolyzes cholesterol ester into cholesterol and cholesterol oxidase converts the enzymatically generated cholesterol to 4-cholesten-3-