Keywords: biosensor, algae, Chlorella, alkaline phosphatase, fluorescence, heavy metals, ecotoxicology

A biosensor is constructed to detect heavy metals from inhibition of alkaline phosphatase (AP) present on the external membrane of Chlorella vulgaris microalgae. The microalgal cells are immobilized on removable membranes placed in front of the tip of an optical fiber bundle inside a homemade microcell. C. vulgaris was cultivated in the laboratory and its alkaline phosphatase activity is strongly inhibited in the presence of heavy metals. This property has been used for the determination of those toxic compounds.

I. Introduction

It is known that heavymeta ls have a toxic effect on living organisms through their action on enzymes of the metabolic chain. The discharge of those compounds into nature leads to the requirement for methods that can monitor them in low concentrations for their control in the environment [START_REF] Gu | A two stage minibioreactor system for continuous toxicity monitoring[END_REF][START_REF] Schubnell | An ISFET-algal (Chlamydomonas) hybrid provides a system for eco-toxicological tests[END_REF]. Heavy metals are usually determined by sophisticated techniques like ionic chromatography, mass spectrometry, inductively coup led plasma, and polarography, which need to be carried out in laboratories after sampling. These methods are not suitable for continuous monitoring in the environment [START_REF] Rogers | Update on environmental biosensors[END_REF]. Heavy metals can also be determined with ion-sensitive electrodes; however, the electrode selectivity and response cannot be associated with the toxicity impact of those chemicals. Therefore, enzyme sensors have been studied and constructed for determination of inhibitors [START_REF] Gayet | Detection of heavymetal salts with biosensors built with an oxygen electrode coupled to various immobilised oxidases and dehydrogenases[END_REF][START_REF] Cowell | The detection and identification of metal and organic pollutants in potable water using enzyme assays suitable for sensor development[END_REF]; they are based essentially on the inhibition of purified enzymes like urease, cholinesterase, and cytochrome oxidase [START_REF] Tran-Minh | Immobilized enzyme probes for determining inhibitors[END_REF][START_REF] Albery | Inhibited enzyme electrodes. Part 2: The kinetics of the cytochrome oxidase system[END_REF]. To provide inexpensive biosensors for environmental control, the enzyme purification step needs to be avoided and enzymes are kept as such in their natural system to ensure a better lifetime [START_REF] Samson | Use of algal fluorescence for determination of phytotoxicity of heavy metals and pesticides as environmental pollutants[END_REF][START_REF] Rechnitz | Biosensors based on cell and tissue material[END_REF]. This approach was adopted with the use of alkaline phosphatase located on the external membrane of the whole cell Chlorella vulgaris for direct determination of heavy metals. On the other hand, this enzyme, which plays a key role in phosphate metabolism, is easily accessible to pollutants since it is attached to the outer part of the cell membrane. In vivo inhibition of this enzyme reflects the natural deletorious effect of toxicants on microalgae and represents a real ecological interest considering that algae are involved in the primary step of the food chain.

II. Material and Methods

I II I. .1 1. . A

Al lg ga al l M Me em mb br ra an ne e P Pr re ep pa ar ra at ti io on n The Chlorella vulgaris strain (CCAP 211/12) was purchased from The Culture Collection of Algae and Protozoa at Cumbria, United Kingdom. The axenic algal strain was grown in the culture medium and under conditions described by the International Organization for Standardization (ISO, 1989). To induce a high alkaline phosphatase concentration in the algal cell, the algal solution was centrifuged and the pellets were resuspended in phosphate-free medium to starve the algal cells for 25 days which corresponds to maximal enzyme activity. This resuspension was carried out under the same conditions and with the same culture medium as that used previously for the growth of this strain except there were no phosphate ions. Although the medium did not contain any phosphate, an increase in cell density was observed, particularly at the beginning of the subculture period, due to the consumption of the phosphate remaining in the cell. This phosphate consumption induces an increase in alkaline phosphatase specific activity. Two milliliters of the resulting aliquot was then filtered on a glass microfiber filter (GF/C Whatman, 45.7-mm filter diameter, 1.2-mm pore diameter). Punching this filter provides small disks (8mm in diameter) which can be fitted to the optical transducer later. A value of 0.73 g/L dry wt was obtained after the filter had been dried for 2 h at 1051C and placed in a desiccator for 0.5 h. I II I. .2 2. . O Op pe er ra at ti in ng g P Pr ri in nc ci ip pl le e Determination of alkaline phosphatase (AP) activity is carried out with methylumbelliferoyl phosphate (MUP) as substrate dissolved at various concentrations in a Tris-HCl (0.1M, pH 8.5) buffer solution containing 1mM MgCl2. This solution is referred to as MUP solution hereafter. MUP was purchased from Sigma; other chemicals from Merck. The reaction product methylumbelliferone (MUF) is fluorescent:

( ) AP MUP MUF fluorescent phosphate  
AP activity can easily be measured from the MUF fluorescence emission (460nm wavelength) under excitation light (350nm wavelength) when the MUP solution is brought into contact with AP. This can be done either by adding the MUP solution into the microplate wells containing C. vulgaris or by injecting the MUP solution into the carrier stream leading to an algal biosensor. In the latter case, an optical fiber is required to send the excitation light to the algal membrane and to collect the fluorescence radiation up to the fluorometer. Since flow injection analysis (FIA) technique is involved, AP activity is determined from peak area (or peak height) corresponding to the concentration of substrate injected. For inhibitor determination, a given amount of heavy metal is introduced into the MUP solution to have the required concentration. A reduction in peak area or peak height is observed with use of the inhibitor solution. I II I. .3 3. . C Co on ns st tr ru uc ct ti io on n o of f t th he e A Al lg ga al l O Op pt ti ic ca al l B Bi io os se en ns so or r The biosensor was constructed with a removable algal membrane placed in a 1-mL homemade flow cell [START_REF] Naessens | Fiber optic biosensor using Chlorella vulgaris for determination of toxic compounds[END_REF]. Opposite the membrane, a bifurcated bundle of randomized optical fibers was oriented so that the incident light hit the upper part of the membrane and the resulting fluorescence could be collected by the corresponding fibers and transmitted to the Spex Fluorolog 2 fluorometer (from Jobin-Yvon) equipped with a microcomputer for data recording. This device was quite suitable for the assay under FIA conditions which require a very small amount of substrate for measurement of enzyme activity.

III. Results

I II II I. .1 1. . I

In nd du uc ct ti io on n o of f A Al lk ka al li in ne e P Ph ho os sp ph ha at ta as se e AP activity in algal cells can be increased by inducing its production with the starvation method. This activity is measured with a batch system under air-saturated conditions. Enzyme activity was plotted as a function of time during the starvation period to assess the maximal specific activity(Figure 1). This sigmoid curve shows that at least 25 days of subculture is required to attain maximal activity. This time was used in our experiments. I II II I. .2 2. . E Ef ff fe ec ct t o of f S Su ub bs st tr ra at te e C Co on nc ce en nt tr ra at ti io on n The configuration of the FIA manifold used is shown in Figure 2. AP activity of algae immobilized on the biosensor is determined with different substrate concentrations. Two substrates can be used: p-nitrophenyl phosphate (pNPP) and MUP. The former produces pnitrophenol, which absorbs at 405 nm, and the latter gives MUF, which fluoresces at 460 nm. The second substrate was chosen for its higher sensitivity and its simplicity in biosensor design. The substrate MUP was placed in a Tris buffer solution (0.1 M, pH 8.5) containing 1mM MgCl2 as enzyme activator, and AP activity was measured in terms of fluorescence intensity from the biosensor response. Various substrate volumes to be injected in the carrier stream were tested: a volume of 0.5mL represents a good compromise to achieve an acceptable response time and good repeatability. A flow rate of 8.4 mL.min-1 was also used for the same reasons. The effect of toxic chemicals on AP was tested with cadmium and lead as heavy metals. Enzyme activity was determined from the fluorescence measurements in the presence and absence of metal in the 0.1mM MUP solution which nearly corresponds to the maximal velocity of the enzyme. When the substrate solution containing the metal is injected, the enzyme is inhibited, and peak area and peak height are reduced. A percentage of enzyme inhibition is introduced to emphasize the biosensor response to heavy metals as inhibitors while reducing experimental variations. It is defined as a ratio of the biosensor response in the presence of a heavy metal to the response in the absence of inhibitor. These responses can be compared with the results obtained with a microplate reader (Figure 4).

IV. Discussion

The responses of the algal biosensor using immobilized alkaline phosphatase to cadmium and lead are in agreement with the results obtained with a microplate reader based on the inhibition of the same enzyme system. This means that the technique can be extended to other heavy metals such as nickel, mercury, copper, zinc, and chromium that inhibit algal alkaline phosphatase. Heavy metals also inhibit algal growth [START_REF] Vasseur | Influence of some experimental factors on metal toxicity to Selenastrum capricornutum[END_REF]) but this process was not taken into account in these measurements because in this enzyme inhibition based method, determination of peak areas of peak heights in the presence and absence of heavy metals was carried out with a single algal membrane and consequently with the same amount of algae. Therefore, only specific AP activity of Chlorella is concerned in the biosensor response. In addition, use of the percentage inhibition provides a reliable response that does not depend on algal strains. It also reduces the effect of long-term enzyme activity loss. Alkaline phosphatase was induced to have maximal activity: this procedure allows the use of a smaller amount of C. vulgaris and the deposition of a thinner layer, which favors a fast response time. However, an optimal enzyme concentration is needed to improve the inhibitors detection limits [START_REF] Tran-Minh | Immobilized enzyme probes for determining inhibitors[END_REF]. The range of toxic concentrations detected with this biosensor makes it more suitable for determination of heavy metals present in waste water rather than surface water. Other experiments conducted in our laboratory have shown that this enzyme sensor is not sensitive to herbicides because AP is not inhibited by these compounds. Therefore, inhibition of algal photosynthesis cannot be assessed with this enzymatic technique. When herbicides are present together with heavy metals, only the latter is detected with this biosensor. The biosensor response also depends on substrate concentration for a given concentration of heavy metal used as inhibitor. High substrate concentrations corresponding to the maximal rate of the enzyme often reduce enzyme dependence on substrate concentration. However, those concentrations cannot be used for enzymes that are inhibited by an excess of substrate, which interferes with the metal detection also based on enzyme inhibition. An optimal substrate concentration should then be determined from the activity-substrate curve. Flow injection analysis associated with a biosensor provides repeatable results along with a low consumption of substrate. The system can easily be automated for the biosensor to operate in continuous monitoring.

V. Conclusion

Biosensors using immobilized enzymes present in whole cells have potential advantages for determination of toxic compounds as AP inhibitors. The enzyme remains in its natural environment which favors long-term stability and reflects the mechanism of toxic inhibition, being therefore of ecological interest. Biosensors based on inhibition of algal alkaline phosphatase are sensitive to heavy metals; they are complements to those based on photosynthesis inhibition designed for herbicide determination [START_REF] Naessens | Fiber optic biosensor using Chlorella vulgaris for determination of toxic compounds[END_REF].

Figure 3

 3 illustrates typical Michaelis-Menten behavior, and an approximate value of 20 mMMUP for the Michaelis constant is obtained from the curve. I II II I. .3 3. . D De et te er rm mi in na at ti io on n o of f H He ea av vy y M Me et ta al ls s

Figures 5 F

 5 Figures


Keywords: Electropolymerization, Conducting polymer, Glucose enzyme electrodes, FIA

A conducting polymer of poly 3,4-ethylenedioxythiophene (PEDT) was used as a matrix for entrapment of enzymes onto a platinum electrode surface in order to construct amperometric biosensors. Glucose oxidase (GOD) was used as an example, and it was entrapped in the polymer during the electrochemical polymerization. Glucose in oxygenated solutions was tested by amperometric measurements at +650 mV (vs. SCE) in a batch system. The influence of several experimental parameters in the electropolymerization process was explored to optimize the analytical performance. The detection limit and sensitivity for this biosensor were 4×10 -5 M and 15.2 mA M -1 cm -2 , respectively. A linear range of response was found from 0.2 to 8 mM of glucose. The response time was 2-5 s. The stability of the electropolymerized films was evaluated in operational conditions. The glucose probe, stored in buffer at 4 °C when not in use, showed a residual activity of 40% after about 1 month. Glucose in synthetic serum was determined under flow injection conditions using an amperometric flow cell. K

SILVIA FABIANO (1) , CANH TRANH MINH (1) , BENOÎT PIRO (2) , LAN ANH DANG (2) , MINH CHAU PHAM (2) , OLIVIER VITTORI (3)

I. Introduction

Amperometric enzyme electrodes have generally been constructed by direct immobilization of oxidoreductases on electrode surfaces. Platinum, gold, graphite or glassy carbon electrodes have been used. They often have produced unstable responses attributed to a poor charge transfer between the electrode and the enzymes. This is due to the fact that biocatalytic sites are usually buried in the inner part of the enzymatic protein [1]. Therefore, natural or artificial redox mediators are added in the soluble [2] or immobilized [3] form to shuttle the electrons between the various redox chemicals involved in the oxidoreductase process [4]. Conducting polymers have also been used in amperometric enzyme electrodes with the intention to couple the electron transfer reaction between enzyme and electrode via the ramified conducting network of the polymer [2, 5 and 6]. They could also have appropriate functional groups to which allowed enzymes can be attached. Finally, many conducting polymers can be deposited electrochemically [7]. In most polymers such as polypyrrole [8, 9, 10, 11, 12, 13, 14 and 15], polyaniline [16, 17 and 18] and polythiophene [19 and 20], redox charges are delocalised over some polymer groups that could facilitate the oxidation-reduction processes. Among the promising family of polythiophenes, poly 3,4-ethylenedioxythiophene (PEDT) shows remarkable stability, provides homogeneous films [21] and can be synthesized electrochemically even in aqueous medium [22]. Oligonucleotides have then been incorporated into PEDT films by electrochemical method [23]. Therefore, we recently used PEDT to covalently attach glucose oxidase (GOD) onto a glassy carbon electrode to obtain an enzyme electrode for glucose sensing [24]. In this paper, this polymer is directly electrogenerated on a platinum electrode to produce thin films for a single-step entrapment of the enzyme. The amperometric biosensor obtained from this technique is designed to be integrated in a flow injection system to determine glucose in real samples. Amperometric flow injection analysis (FIA) associated with immobilized enzyme sensor has become recognized as a selective, simple, accurate and rapid technique for the determination of a variety of biologically important compounds [25]. In such a system, solid electrode-based detectors are commonly used for the detection of electroactive products generated enzymatically. This system offers advantages of highly reproducible timing and high sample throughput and eliminates the need to attain steady states [26]. As the PEDT biosensor has never been tested in FIA conditions, the PEDT biosensor described in this paper is therefore integrated in a FIA equipment to determine the performance of this coupling.

In order to construct the amperometric enzyme sensor, glucose oxidase is used as an example of a redox protein. The enzyme catalyses, in the presence of molecular oxygen, the oxidation of β-D -glucose into gluconic acid and hydrogen peroxide. The conversion of β-D -glucose to gluconic acid involves the transfer of two protons and two electrons from the substrate to the flavin moiety of the enzyme [27]. The electron transfer from the redox cofactor to the sensing electrode might also be facilitated by the presence of a polymeric conducting material.

II. Experimental

I II I. .1 1. . C Ch he em mi ic ca al ls s 3,4-Ethylenedioxythiophene monomer (EDT) was kindly provided by Bayer (France). Polyethylene glycol (PEG) of 15,000 MW was purchased from Aldrich. -(+)-Glucose and glucose oxidase (EC 1.1.3.4.) type VII from Aspergillus niger were purchased from Sigma (France) and used as received. Accutrol™ normal chemistry control serum was purchased from Sigma. All other chemicals were of analytical grade. Glucose solutions were allowed to mutarotate overnight at room temperature before use. Phosphate buffer was prepared using dihydrogen phosphate dihydrate (Prolabo, France). pH was adjusted with sodium hydroxide solution. All solutions were prepared using milliQ® pure water (Millipore, France). I II I. .2 2. . A Ap pp pa ar ra at tu us s Voltammetric and amperometric measurements were performed using a Tacussel PGP 201 potentiostat. The potentiostat output was recorded and processed by means of a Volta Master 1 software. The electrochemical cell was a three-electrode cell where the PEDT/GOD-modified platinum disk electrode acted as the working electrode and a platinum wire as the counter electrode. All potentials were measured vs. SCE. The working electrodes were constructed using a BAS® platinum electrode with a disk-shaped active surface with a diameter of 1.6 mm for batch system and a diameter of 3 mm for flow injection system. A miniaturized flow cell from BAS® was used for FIA measurements (Figure 1). One side of the flow cell contains the inlet and outlet cannulas, the auxiliary electrode and the reference electrode. The other side holds the working electrode. The reference electrode is an electrode without porous junction. It consists of silver metal (Ag) coated with a layer of silver chloride salt (AgCl). The auxiliary electrode is an adjacent 2-mm stainless-steel disk. The inlet cannula is connected to the source of the flow stream, and the outlet cannula is directed to a waste container. The flow system volume using a 51-μm-thick Teflon gasket is 6.9 μl.

All the lines were made from a 0.41-mm internal diameter Teflon tubing. A sample line was used to fill the sample loop on a Rheodyne (USA) sample injection valve that was used to inject samples into a buffer stream. An Ismatec peristaltic pump (Switzerland) was used in conjunction with an injection valve.

I II I. .3 3. . I

Im mm mo ob bi il li iz za at ti io on n p pr ro oc ce ed du ur re e Bare electrodes used for preparing the biosensors consisted of Pt disks. The working electrode surface was polished before use with alumina (polishing alumina fluid provided by BAS®, England) of 0.05 μm particle size. After careful rinsing with distilled water, the electrodes were pretreated by potential cycling in 0.5 M H 2 SO 4 from -1500 to +1500 mV (vs. SCE) at a scan rate of 2 V/s until no changes were observed in the cyclic voltammograms. A film of enzymatic PEDT was deposited on the Pt electrode by electropolymerizing 3,4ethylenedioxythiophene (EDT) through cyclic voltammetry. A 10 -2 M solution of EDT was obtained by dissolving EDT in a 0.02-M, pH 6.2 phosphate buffer solution containing 10 -3 M PEG and 1000 IU glucose oxidase ml -1 . Electrochemical growth of PEDT was carried out using potential cycling between +200 and +1500 mV at a scan rate of 0.1 V/s for a total of 15 cycles. The same immobilization procedures were used to construct glucose enzyme sensor for batch and flow injection systems. I II I. .4 4. . B Ba at tc ch h m me ea as su ur re em me en nt ts s For the batch amperometric measurement of glucose, the PEDT/GOD biosensor was dipped in a 20-ml beaker containing 10 ml of potassium phosphate buffer solution. The electrochemical cell was completed with a SCE reference electrode and a platinum wire counter electrode. The solution was continuously stirred in the open air with a magnetic stirrer to maintain a constant oxygen concentration in the solution for the enzymatic reaction to operate properly. The temperature was controlled using a thermostat (Lauda-Thermostat, Germany). After the cell was assembled, a potential of +650 mV was applied to the working electrode to detect hydrogen peroxide. When a baseline was reached, glucose was spiked. The current generated on the biosensor was measured when a steady state was reached. Unless otherwise stated, all experiments were carried out at 25 °C, with an applied potential of +650 mV vs. SCE, in a sodium phosphate buffer (0.01 M, pH 7.0) containing 0.1 M KCl. I II I. .5 5. . F Fl lo ow w i in nj je ec ct ti io on n m me ea as su ur re em me en nt ts s The configuration of the FIA manifold has already been described [28]. A peristaltic pump was used to transport the carrier liquid and inject the sample and an injection valve with a loop to receive sample and carrier liquid alternately. The determination of glucose was obtained with a biosensor placed in a miniaturized flow cell. An potential of +650 mV vs. Ag/AgCl was applied to the amperometric enzyme electrode by means of a potentiostat connected to a recorder. In this method, time, temperature and dispersion were all reproducible conditions. The peristaltic pump operated with a 75 μl min -1 flow rate of a carrier stream of potassium phosphate buffer (0.01 M, pH 7.0) containing 0.1 M KCl. The volume of injection loop was 50 μl. The response curve appears in the form of peaks whose heights depend on the substrate concentration. The concentration of the analyte in a sample can thus be determined from a calibration curve setup with known substrate standards in duplicate measurements.

III. Results and discussion

I II II I. .1 1. . O
Op pt ti im mi iz za at ti io on n o of f t th he e e en nz zy ym me e i im mm mo ob bi il li iz za at ti io on n Current-potential curves at steady state for glucose are shown in Figure 2. The current response increases as soon as the potential is over +200 mV and a plateau is reached at +500 mV. Therefore, +650 mV was chosen as the working potential for glucose determination. It corresponds to the oxidation of H 2 O 2 produced by the enzyme-substrate reaction. To control the electroentrapment of the enzyme, electropolymerization of this solution was carried out using a number of voltammetric cycles: The first cycles were applied to induce the polymerization process and the following cycles to achieve the overall coating of the electrode. Fifteen cycles were found sufficient to ensure an effective enzyme immobilization, too many cycles would have deleterious effects on mechanical properties of the polymer layer [23]. The thickness of the polymer film can be controlled by the charge consumed in the polymerization of the monomer. The enzyme concentration has little effect on the membrane thickness because the matrix is mainly represented by PEDT [2]. The influence of enzyme loading in the polymer on the sensor response to glucose was examined (Figure 3). Responses of the PEDT/GOD biosensor to 5 mM glucose increase when the enzyme concentration in the electropolymerization solution increases from 50 to 800 IU ml -1 and reach a plateau at 800 IU ml -1 corresponding to the enzyme saturation in the solution. However, a 1000 IU ml -1 concentration, which is the saturation limit, was adopted for enzyme immobilization in case of enzyme activity loss during electropolymerization. A specific activity of immobilized GOD is estimated to be 0.4 IU/ml from spectrophotometric determination of glucose oxidase activity described elsewhere. I II II I. .2 2. . R Re es sp po on ns se e o of f t th he e b bi io os se en ns so or r i in n a a b ba at tc ch h s sy ys st te em m Figure 4 shows steady-state amperometric responses of the PEDT/GOD biosensor to the addition of aliquots of the stock glucose solution into an air-saturated phosphate buffer solution in a batch system. Oxygen concentration should be kept constant because it is a cosubstrate of the enzymatic reaction. These responses are used to construct a calibration curve for glucose. Although a time lag is necessary for the sample solution to be homogeneous after substrate addition, a rapid response of the PEDT/GOD biosensor is observed for each glucose concentration: 2-5 s is estimated to reach a stable plateau. The enzyme electrode calibration curve using a PEDT film is presented in Figure 5. The data are obtained from measurements repeated three times for each concentration, and the mean value is reported in this figure. The linear range of the biosensor response to glucose concentration from 0.2 up to 8 mM would allow to determine human blood glucose, which lies within the narrow limits of 3.5-5.0 mM. The detection limit of the enzyme electrode is 40 μM, which corresponds to a signal-to-noise ratio of 3. The sensitivity corresponding to the linear range is about 15.2 mA M -1 cm -2 . This sensitivity gradually decreases at higher glucose concentrations. I II II I. . 3

. E

Ef ff fe ec ct ts s o of f p pH H The effect of pH of the electropolymerization solution on the amperometric response of the PEDT/GOD electrode was studied for pHs ranging from 5 to 8 (Figure 6). Optimum response to 5 mM of glucose was observed for an electropolymerization at pH 6.2. This pH value was previously adopted for polymerization of the same polymer [21]. An optimized polymerization pH should allow an efficient entrapment of the enzyme while preventing loss of enzyme activity under polymerization conditions. PEG was added to the EDT solution to improve the enzyme activity by increasing the hydrophicity of the polymer. The enzyme sensor response also depends on the working pH of the sampling solution. The effect of pH on the behaviour of the enzyme electrode was studied with 0.01 M phosphate buffer solution containing 5 mM glucose. The pH of the solution ranges from 5 to 8. The response of biosensor as a function of working pH is shown in Figure 6. A maximal response is obtained for pH 7.0, which is different from the optimal electropolymerization pH and higher than the optimal pH of glucose oxidase dissolved in a solution which has been reported as 5.6 [START_REF] Barman | Enzyme Handbook[END_REF]. In fact, the kinetics of immobilized enzyme is different from the enzyme kinetics in a homogeneous solution. On the biosensor where the diffusion-reaction process should be taken into account, the oxidation of glucose by immobilized glucose oxidase in the active layer produces gluconic acid, which decreases the local value of pH. The pH value in the enzyme layer bound to the electrode then is lower than the pH of the bulk solution. In order to have the enzyme optimal pH in the PEDT film containing GOD, a higher pH is necessary in the sample solution. I II II I. .4 4. . S St ta ab bi il li it ty y Thermal stability of the enzyme electrode was investigated up to 50 °C. The response of the PEDT/GOD electrode to glucose increases with temperature from 15 to 30 °C, then a decrease in enzyme activity is observed (Figure 7). A 25 °C temperature is adopted in our experiments in order to prevent any risk of enzyme denaturation.

The main objective of enzyme immobilization on a transducer for analytical purposes is to stabilise the enzyme for the biosensor to be used repeatedly over a long period. The long-term stability of a biosensor is then tested everyday with the PEDT/GOD electrode. The biosensor was stored between measurements in a phosphate buffer solution at 4 °C. This testing procedure is carried out with a 5-mM glucose solution over 4 weeks. The enzyme electrode retained about 40% of its initial activity after 25 days (Figure 8). I II II I. .5 5. . F Fl lo ow w i in nj je ec ct ti io on n s sy ys st te em m f fo or r t th he e d de et te er rm mi in na at ti io on n o of f g gl lu uc co os se e i in n s sy yn nt th he et ti ic c s se er ru um m s sa am mp pl le e The batch system previously described is quite suitable to study the effects of pH and temperature on the response of the biosensor as well as its long-term stability. In such a system, the various parameters can easily be adjusted and controlled in order to allow normal working conditions of the biosensor to be optimized. In biochemical analysis, the need of high sample throughput requires continuous measurements and limitation of manual handling. FIA associated with a biosensor provides a fast, reliable and continuous method to determine glucose without any additional reagent. Repeatability of FIA measurements is conducted with injections of glucose in duplicate at various concentrations (Figure 9). A mean RSD value of 3.7% is obtained from those responses. Figure 10 shows typical calibration graph for glucose. A linear relationship between the signal and the concentration is observed up to 10 mM for glucose. The sensitivity calculated from the linear range is about 7.8 mA M -1 cm -2 . The detection limit with this FIA biosensor is 5 μM for a glucose concentration corresponding to a signal-to-noise ratio of 3. This good detection limit is due to a very stable baseline: It is eight times lower than that obtained from the batch system. Monitoring of glucose in blood is important for diagnosis and surveillance of diabetes, the assays should also be conducted with glucose in serum samples to assess the practical applicability of this method. In this work, the control human serum used consists of an assayed lyophilized preparation containing the analytes commonly observed in man following physiological or pharmacological intakes. Some electroactive species except ascorbic acid are present in this control serum at their usual concentrations: acetaminophen (86-139 μM), salicylate (0.5-0.8 mM) and uric acid (321-393 μM). Table 1 shows the results obtained from this FIA biosensor in comparison with others techniques used to determine glucose in human control serum. Assigned means are of replicate assays performed in reference laboratories using the indicated procedure. The concentrations obtained with this FIA biosensor are similar to those given by other procedures.

IV. Conclusion

In this study, the possibility of using conducting polymer as a support for single-step immobilization of glucose oxidase has been investigated. The use of PEDT as an entrapment support of glucose oxidase in the development of a glucose biosensor is motivated by the conductivity and stability properties of PEDT [21, 30, 31 and 32]. We have shown that electrochemical polymerization of EDT results in an efficient technique to immobilize glucose oxidase onto a platinum electrode surface. The batch system provides an easy method to optimize the various operating parameters of the biosensor. Its rapid response time and stability are compatible with its application in flow analysis systems. The determination of human blood glucose is possible due to its wide dynamic range. In spite of a large number of species present in control human serum, the results given by the FIA/biosensor are in agreement with those obtained from other methods currently in use. This immobilization technique, which can be extended to other oxidoreductases, is quite suitable for the design and construction of microbiosensors.

FiguresFFF

  Figures

  In nt te er rn na at ti io on na al l U Un ni it ts s. .

	Tables
	T Ta ab bl le e 1 1: : C Co om mp pa ar ri is so on n o of f t th he e r re es su ul lt ts s o ob bt ta ai in ne ed d w wi it th h t th hi is s F FI IA A/ /b bi io os se en ns so or r t to o t th ho os se e o ob bt ta ai in ne ed d w wi it th h s st ta an nd da ar rd ds s g gl lu uc co os se e k ki it ts s, , i in n s se er ru um m g gl lu uc co os se e a an na al ly ys si is s
	M Me et th ho od d	S SI IU U a a u un ni it t	S SI IU U M Me ea an n	S SI IU U R Ra an ng ge e
	Abbott Spectrum/VP series (hexokinase)	mmol/l	4.9	4.4-5.4
	Bayer-Opera/RA series (hexokinase/340 nm)	mmol/l	5.0	4.5-5.5
	Beckman-CX/LX/Synchron series (hexokinase)	mmol/l	4.8	4.3-5.3
	BioChem-ATAC (hexokinase/340 nm)	mmol/l	5.3	4.7-5.9
	Ortho-Vitros (oxidase)	mmol/l	4.8	4.3-5.3
	Pointe Scientific (oxidase)	mmol/l	5.0	4.5-5.5
	Roche-Cobas Mira series (hexokinase/340 nm)	mmol/l	4.6	4.2-5.0
	Sigma (enzymatic, Trinder)	mmol/l	4.7	3.7-5.7
	Present FIA biosensor	mmol/l	4.1	3.9-4.3
	a a S Sy ys st te em m o of f I