METHOD AND APPARATUS FOR CHEMICAL ANALYSIS

Abstract: A method and an apparatus for chemical analysis of samples of varying matrices is disclosed. The samples are transported in series by a carrier fluid. At a test location is provided a fluid elemement the chemical composition of which is maintained essentially constant in consequence of the exposure of the fluid element to the carrier fluid prior to and subsequent to the exposure of the fluid element to each individual sample.

Description: METHOD AND APPARATUS FOR CHEMICAL ANALYSIS The present invention relates to methods of determining the concentration of an analyte in each individual sample of a sequence of samples by means of an analyte sensor. The methods according to the invention has the features of claim 1 or claim 2 or claim 3. The preferred embodiments have the features of claims 4-9. The analyte sensor may be an electrochemical sensor, an optical sensor or any other suitable sensor means. The fluid element is stationary and is held by any suitable support means, such as a porous or fibrous member in the shape of a film, a sheet, a block, a cylinder or any other suitable shape, e.g. a ceramic pin or a thread, a hydrophilic or porous membrane, a gelled structure etc.. The chemical composition of the fluid element may be maintained by conditioning the fluid element with the carrier fluid prior to the contact with each individual sample. The fluid element is preferably an aqueous composition. The analyzer according to the present invention has the features of claim 10 and the preferred embodiments have the features of claims 11-13. Generally, the analyzer according to the invention is an analyzer with a sample inlet, a sample conduit, a sensor means arranged proximate to a test location of the sample conduit and a sample outlet and further with support means for the fluid element arranged between the actual fluid sample under test and the sensor means. The sample may be fed to the analyzer in any conventional manner, e.g. by injection or aspiration. The sample is transported through the analyzer by $ means of a flowing carrier fluid. The carrier fluid may constitute the conditioning medium of controlled chemical composition. There may be direct contact between the sample and the carrier fluid or there may be a plug of an immiscible fluid such as a plug of oil or of gas separating sample and carrier fluid. In the first instance the analyzer may be of the conventional flow injection analyzer type. Normally, the analyzer will be able to handle in series more than 30 samples per hour. Numerous articles, patents and other publications describe general features of analyte sensors and analyzers. The contents of these publications is incorporated herein by reference. The invention has its primary application in connection with analysis procedures, wherein the individual samples of a sample sequence have differing matrices. E.g. this will be the case with samples of sewage, samples from fermentation plants and samples of whole blood, as long as no sample preparation such as a dilution is carried out prior to the test. The invention will be further illustrated by means of the drawing in which Fig. 1(a) shows a FIA manifold for the enzymatic determination of urea; o Fig. 1(b) shows an exploded view of a gas diffusion/detection unit shown in Fig. 1(a). Fig. 2 shows a calibration graph for urea? Fig. 3 shows a FIA manifold for the detection of wea using an ion-selective electrode; 5 Fig. 4 shows a FIA manifold for the detection of glucose or lactate; Fig. 5 shows calibration curves for glucose; Fig. 6 shows a 25 mM glucose standard record; Fig. 7 shows the glucose response for a glucose o standard with and without paracetamol. In further detail Fig. 1 (a) shows a FIA manifold for the enzymatic determination of urea, comprising an integrated gas diffusion/detection unit, an exploded view of which is shown in (b) . The sample (S) is injected into the donor stream (buffer) and propelled to part B of the flow cell, in which urea is degraded in the enzyme-containing membrane (M) . The generated ammonia diffuses via the gas permeable membrane (G) into the acceptor (indicator) stream, where the resulting pH change is monitored by reflectance measurement by an optical fibre (F), bifurcated at the remote end to communicate with a light source and a photo diode detector (D). I and O denote in- and outlets for the donor and acceptor streams, respectively; GF, gold foil; W, waste. The flow cell, which is made part of a FIA microconduit, is assembled by means of 4 screws. Fig. 2 shows a calibration graph for urea in the range 0-35 mM as obtained with the FIA system depicted in Fig. 1. Fig. 3 shows a FIA manifold for -the detection of o urea using an ion-selective electrode with detection of enzymatically generated ammonium, comprising a syringe pump (PI); a peristaltic pump (P2); an injection valve; a pH meter; a urea electrode (E2) ; and a flow-through reference 5 electrode (El). S - sample injection; FC - flow cell; and - waste. Fig. 4 shows a FIA manifold for the detection of glucose or lactate using an amperometric electrode with detection of enzymatically 0 generated hydrogen peroxide, comprising a syringe pump (PI); a peristaltic pump (P2); an injection valve; a BBC computer; a potentiostat; and a glucose electrode (GE) or a lactate electrode (LE) . S - sample injection, FC - flow cell; and W - waste. 5 Fig. 5 shows calibration curves for glucose in the concentration range 0-40 mM at three different flow rates (0.50, 0.75 and 1.00 ml/min) using the FIA system depicted in Fig. 4. Fig. 6 shows a 25 mM glucose standard recorded at io high paper speed to demonstrate the fast rate of reaction. Fig. 7 shows the glucose response for a 15 mM glucose standard with and without 360 uM paracetamol as a function of the applied is potential. The invention will now be further illustrated- by means of the following examples. EXAMPLE 1 DETERMINATION OF UREA IN UNDILUTED BLOOD SAMPLES 5 BY FLOW INJECTION ANALYSIS USING OPTOSENSING ABSTRACT A miniaturized flow injection system for the assay of urea in undiluted whole blood is described. Based on the optical determination of o ammonia, the system incorporates a flow cell combining gas diffusion and optosensing, the separating barrier betweeen the donor and accepting streams consisting of a sandwich of a hydrophilic membrane between which is contained a s gel of covalently immobilized urease. The sample urea content is quantified by the colour change of an acid-base indicator contained within the acceptor stream. Within the physiological range the measurement is not affected by variations in 0 the pH value, the buffering capacity or the . hematocrit level of the sample solution, nor is it, due to the geometric construction of the flow cell, prone to interferences owing to the colour of the sample solution itself. The system is 5 stable for at least one week of continuous use. INTRODUCTION Urea is the final product of human protein metabolism. Determination of blood urea makes it posible to trace and follow changes in protein ° and amino acid composition, and the concentration of urea in blood is an important index of renal, function. In clinical chemistry blood urea is traditionally asayed in the serum or plasma fraction, yet the ultimate goal is to perform the measurement in the original sample matrix since this requires a minimum of manipulations, and hence a shorter time to obtain the analytical result, such an apparoach furthermore facilitates the possibility of continuous monitoring. Besides, the importance to confine manual handling of blood samples has been emphasised in recent years due to increased occurrence of contagious diseases. Blood urea is almost exlusively assayed via enzymatic degradation by urease CO(NH 2 ) 2 + H 2 0 + H 3 0 + UreaS 1 2NH* + HC0~ followed either by direct determination of ammonium or carbonate, or via indirect procedures where one of these species is determined by means of chemical derivatization, or where the gaseous form of one of the two protolytes, or the ensuing pH change caused by it, is measured. Although urease is a rather inexpensive enzyme, procedures employing immobilized rather than soluble enzymes are preferred, because immobilization of the catalyst not only offers increased stability of the activity of the enzyme, but readily permits its integration into automated systems. Immobilization of urease can be -achieved in many ways (1-4), and the majority of these techniques provide excellent reactors and/or membranes which are stable for extended periods of time. Urea is frequently determined by the indophenol colorimetric method (5), yet most of the work reported in the literature has been devoted to development and use of enzyme electrodes. The nonactin/monactin electrode (2) provides poor o selectivity for ammonium ions over potassium, but the interference can be compensated for by introducing a second ammonium selective electrode for NH- and K background detection (4). Sensing of the pH changes generated in a 5 urease-gel, wrapped around a pH electrode, is another approach (6-8). This type of ammonia sensor yields excellent selectivity for ammonia over ions in the solution (9), yet besides being pH sensitive its response is rather slow. A o hybride urea sensor consisting of an immobilized layer of nitrifying bacteria with ultimate detection of oxygen with an oxygen electrode has been reported (10). An ammonia sensitive In/Pd MOS capacitor (11-13) has been suggested as an s alternative to the ammonia gas electrode with the added advantage of being inexpensive. In the present paper a FIA microconduit system for the assay of urea in undiluted whole blood is described. Based on the optical determination of 0 ammonia, the system incorporates a flow cell combining gas diffusion and optosensing, the separating barrier between the donor and accepting streams consisting of a sandwich of a hydrophobic gas permeable membrane and a hydrophilic membrane between which is contained a gel of covalently immobilized urease. The ammonia generated by the enzymatic degradation of the urea in the sample solution is determined via the colour change of an acid-base indicator contained within the acceptor stream, o the change in reflectance being a function of the ammonia diffused through the gas permeable membrane and thus the sample urea content. Within the physiological range the measurement is not affected by variations in the pH value or the 5 buffering capacity of the sample solution, nor is it, due to the geometric construction of the flow cell, prone to interferences owing to the colour of the sample solution itself. EXPERIMENTAL ° Apparatus All measurements and experiments were executed in the FIA manifold depicted in Fig. la, comprising an integrated gas diffusion and detection unit (for details, see below) accommodating an 5 acceptor (indicator) stream and a donor (buffer) stream. The sample solution (20μl) is injected into the latter, the distance between the injection port and the the gas diffusion cell being made as short as possible (2,5 cm, corresponding to 5 μl ) . The rotary injection valve (15) and the gas diffusion/detection unit were incorporated into a FIA microconduit (14, 15). The gas diffusion/detection unit (Fig. lb) was constructed of black PVC. The unit comprises two separate parts (A and B) between which, when assembled by four screws, the membrane sandwich is contained. Into part A is milled a 1 x 1 x 20 mm groove communicating with two perpendicularly drilled holes (0,5 mm ID) serving as in- (I) and outlets (0) for the acceptor stream. Approximately 5 mm from the outlet a perpendicular hole is drilled to accommodate an optical fibre (Eska Extra, EM 4001, 1,0 mm OD fibre, Mitubishi Rayon Co.). Part B embodies a 1 x 1 x 10 mm groove served by in- and outlets for the donor stream. The membrane sandwich consists of a gas permeable membrane (G) <Celgard 2500, hydrophobic polypropylene polymer, thickness 25 μm, pore size 0.004 ^i , Celanese, Belgium), and an enzyme membrane (M) (see below). The optical measurement of the colour of the acid-base indicator is effected by reflectance, the nontransparent white gas permeable membrane serving as a reflecting opaque background, which, in order to further improve the reflected signal, is supplemented by the incorporaration of a thin foil of gold (GF) placed opposite the optical fibre. Note, that the sample outlet for the donor (buffer) stream is located before the position of the optical fibre thereby obliviating any interference due to colour of the sample stream itself. The donor and the acceptor streams were propelled by means of an Ismatec Mini-S-840 peristaltic pump at 0.1 ml/min. and 0.3 ml/min., respectively. The colour change in the acceptor stream was monitored by a home-made fibre optical diode photometer which transmits and receives through a single optical fibre. The photometer is constructed to emit chopped light at two different wavelengths, that is, at 820 nm, at which the absorbance of the acid-base indicator (bromothymol blue, BTB) is independent of pH (i.e., the ensuing reflected signal received at this wavelength serves as a reference level), and at 635 nm, where the absorbance of BTB is a function of pH. Since the emitted light is chopped, interference due to possible changes in the stray light jLevel is automatically eliminated. The signal from the optical fibre photometer was sent to a recorder (Radiometer Servograph REC 61, furnished with a REA-112 high-sensitivity interface), the peak height at peak maximum serving as the analytical readout. Reagents All chemicals were of analytical-reagent grade, and redistilled water was used throughout. The carrier (donor) stream solution consisted of 0.1 M TRIS buffer adjusted to pH 9.2. The acceptor stream solution was 0.04% (w/v bromothymol blue (Merck) dissolved in 24% aqueous ethanol. Prior to use the pH of the solution was adjusted to 6.2-6.3 with 1.0 M NaOH, s corresponding to 5-10% of the indicator initially being on the alkaline form, the blue colour of which was monitored at 635 nm. The aqueous urea standards in the range 1-35.0 mM used for calibrating the system were prepared by o appropriate dilutions with 0.01 M phosphate buffer of pH 7.2 of a 35.0 mM stock solution made by dissolving 2.102 g of urea (Riedel-de Haen) in 1 1 of 0.01 M phosphate buffer of pH 7.2. Human blood and serum samples were obtained from s ,the clinical chemistry department of a local hospital (Copenhagen County Hospital in Herlev) . Preparation of Immobilized Enzyme Membranes 40 mg of urease from "jack beans" (E.C. 3.5.1.5, Sigma, 70 U/mg) and 30 mg of bovine serum o albumine were completely dissolved in 0.3 ml of phosphate buffer (10 mM, pH 7.2) and 20 11 of glutaraldehyde (25% aqueous solution) was added. The solution became viscous within 2 min., whereupon 20 jul of this mixture was placed on 8.5 2 s mm of a hydrophilic membrane (Nucleopore, pc black, thickness 10 um, pore size 0.2 urn) and covered with a similar hydrophilic membrane to protect the gel of the covalently immobilized urease. When not in use the membrane sandwich was stored in a refrigerator (4°C) where it is stable for several months without any detectable deterioration. RESULTS AND DISCUSSION s Preliminary Experiments In order to determine the optimum operational parameters, a series of preliminary investigations were conducted to determine the influence on the analytical value of (a) the o hematocrit value (i.e., the fraction of blood cells), and (b) the buffering capacity of the sample solutions. A controlled dispersion of the sample zone while it is passing through the system towards the s detector is essential in Flow Injection Analysis. Betteridge et al. (16) have shown that the viscosity of the sample will influenc the dispersion process, and since the hematocrit level - and hence the viscosity - in blood samples from different individuals vary to a certain extent, the influence of this parameter .was initially examined. This was executed by means of a single-line FIA manifold furnished with an optical flow through -cell (18 jul) placed in a photometer between which and the injection valve variable lengths of coils could be inserted (0.82 mm ID, of lengths 10, 20, 30 and 40 cm, respectively). A series of plasma samples containing 0, 25, 50 and 75% blood cells were prepared and injected (20,ul) into the FIA system where the carrier stream (isotonic buffer, pH 7.2) was pumped at 0.1 ml/min. The optical signal was measured at 620 nm. The extent of the relative dispersion was evaluated by measuring the widths of the peaks at the base, the experiments being randomized. The results did indeed show that the dispersion was profoundly influenced by the hematocrit level. Thus, not being able to control the dispersion of the blood samples with sufficient accuracy at dispersion " coefficient (D) values above 2, the FIA manifold should therefore for practical assays be operated at a D-value as close as possible to 1 i.e. less than 1.1, preferably less than 1.05, (where the influence of changes in the hematocrit level is negligible, cf. the discussion below), that is, the distance between the injection port and the detection position should be minimized. Per definition D= ■_= , Co being the concentration of the species under test in the sample before injection in the carrier stream and C being the maximum concentration of the species under test provided in the sample zone of the carrier stream at the detection position. The pH of the carrier (donor) solution was chosen as a compromise between the optimum pH of the enzyme activity (pH 7-8) and the degree of conversion of the enzymatically generated ammonium to ammonia (pK 9.25). In order to reach a sufficient degree of conversion without losing enzyme activity, a pH of the carrier stream of 9.2 was selected. However, since blood samples, which normally lie in the pH range s 7.2-7.6, have variable buffering capacities, it was investigated whether this variation exerted any influence on the measured signal. A series of urea standards in phosphate buffers of different pH values (7.2-7.6) and different o buffering capacities (/2> = 0.003-0.01) were prepared and injected into the system depicted in Fig. la. Irrespective of pH and buffering capacity, practically the same readout for the individual urea standards was found, the s variation within the individual series being of the same order of magnitude as between the different series (of the order of 1-2%). Thus, it can be concluded that any influence of response due to variations in buffering capacity o and pH in human blood samples can be neglected under thjs conditions prevailing in the FIA system. In order to dertermine optimum flow rates in the FIA system used for the actual assays, various flow rates of the donor and acceptor streams, and the ratio between them, were examined. Since the time of exposure of the sample to the enzyme will determine the extent of substrate conversion - and hence the amount of ammonia diffused into the acceptor stream - and since the pK value of the acid-base indicator and the buffering capacity of the acceptor stream will determine the pH change obtained for a given amount of ammonia diffused into it, the dynamic measuring range will be a function of the pumping rates. Thus, the pumping rates used in Fig. 1 were chosen to fit a dynamic measuring range of 1-35 s mM urea, corresponding to the physiological range. Sensing Mechanism of the Membrane It might appear that there is a discrepancy between the statement that the optimum experimental conditions found require a D-value i o as close as possible to 1, while at the same time it is specified that the composition of the carrier stream (pH and buffering capacity) is important for a successful assay, because at D = 1 no physical mixing of the carrier stream and 15 the sample solution takes place, and therefore the composition of the carrier stream should have no influence. It is, however, important to realize that a perfectly controlled spatial and- kinetic discrimination prevails in the sandwich - ' membrane operated under FIA conditions, so that the sensing process, in fact, occurs in two steps: In the first step, where the injected sample is in contact with the enzyme layer, the degradation of urea to ammonium takes place. In 25 the next step, when the undiluted sample zone is swept away by the more alkaline carrier stream, the generated and entrapped ammonium ions are being converted to ammonia which is then transferred across the gas permeable membrane 3o into the acceptor stream of the acid-base indicator. Thus, any influence of the matrix of the sample (such as the hematocrit level of blood) is eliminated by performing the enzymatic reaction on the undiluted sample material (D = 1), while the conversion of analyte to a measurable species then is effected through a judicious choice of the carrier stream. Therefore a change in the pH of the carrier stream within a pH range compatible with the enzyme (i.e., where the activity of the immobillized enzyme within the membrane can be considered to remain constant) should almost exclusively reflect the conversion of ammonium to ammonia. This was experimentally verified by varying the pH of the carrier stream from 7.0-9.5, the low pH values yielding lower signals, while the higher pH values resulted in comparatively higher peaks. Harrow and Janata (17, 18), when assaying pH and the carbon dioxide content in whole blood, were the first to point out the difficulties associated with the varying hematocrit level in blood samples from different individuals, observing that the measurements were unreliable unless minimum sample dispersion in the FIA system was guaranteed. In their papers they concluded "that it is interesting to note that if the distance between the detector and the point of injection is minimized, the dispersion approaches unity, in which case the viscosity and hematocrit effect disappear altogether. In that case, however, the flow injection system operates essentially in a steady-state mode (D = 1), and the power to do chemistry by controlled mixing of sample with reagents is largely lost". However, as demonstrated herein it is indeed 5 possible to take advantage of chemical reactions and/or conversions in FIA systems operated at limited dispersion, namely by exploiting the mechamism of the time/space kinetic discrimination. Furthermore, this approach o readily allows optimization of the level of response of a given substrate/ immobilized-enzyme system, a provision being, of course, that the enzyme is compatible with the conditions of the carrier stream solution to which it is exposed. 5 The limit of the alkalinity of the carrier stream for the urea system, adjusted to favour the conversion of ammonium to ammonia, is therefore dictated by the stability of urease in this type of medium, the activity of the enzyme decreasing o rapidly at higher pH values. Since the level of urea in human blood is in the millimolar range, the amount of ammonia ultimately diffusing into the acceptor stream of the system in Fig. 1 is of such magnitude that 5 the ensuing absorbance change registered by the optical system is sufficient to obtain a reliable signal. However, should the conversion efficiency and/or the substrate concentration be lower, additional signal enhancement may be o obtained by operating the donor stream and especially the acceptor stream in a stopped-flow mode (14) . Assay of Standards, Plasma and Whole Blood Samples For actual urea assays the FIA manifold depicted in Fig. la was used. A typical calibration graph obtained with a series of aqueous urea standards is shown in Fig. 2. The dynamic measuring range is 0-35 mM urea with a linear range from 0 to 6 mM urea (the normal physiological range). The curvature of the calibration plot might be due to several factors. In order to obtain a linear relationship between the urea concentration and the recorded absorbance signal the following two conditions must be fulfilled: (a) The kinetics of theι urease enzyme has to follow that of pseudo-first order reaction conditions. (b) The change of pH in the acceptor stream due to diffusion of ammonia must fall within the range pK. * 0.5. These two parameters were not attempted optimized because the main objective of this investigation was to demonstrate the feasability of measuring in whole blood samples. However, it has ben shown by Petersson (19) that urease-containing membranes, prepared according to the procedure described in this communication, readily can be made to contain higher activities, that is, the enzyme reaction will be strictly that of pseudo-first order up to urea concentrations as high as 30 mM. With respect to extending the linear range of the acid-base indicator solution, Ramsing et al. (20) have shown that this can be achieved by using a solution containing several indicators of appropriately spaced pK-values. The standard deviation of the aqueous standards was found to be less than 2% (n = 160) with a sampling frequency of 30 samples/h. Next, a series of 16 blood plasma samples, supplied by the Copenhagen County Hospital in Herlev (CCHH), were assayed in order to compare the FIA method with that used routinely at CCHH where urea exclusively is analysed in plasma. The results are reproduced in Table 1, the correlation between CCHH and FIA being CCHH = 1.09 FIA - 1.33 mM, with a correlation coefficient of 0.993. A comparison of urea measurements on whole blood and plasma (n = 16) was then made. The results showed that the concentration of urea in whole blood was 5-8% less than that of the corresponding plasma samples. This deviation can be ascribed to small variations in the D-value of the FIA system operated with plasma and blood samples, respectively. Further miniaturization of the system (i.e., realization of a D-value as close to unity as possible) has shown that this discrepancy can be eliminated (19). Individual variations in the hematocrit level had no or very little effect «2%) on the results. The reproducibility of measurement on the whole blood samples was comparable to that obtained with the plasma samples/aqueous standards. No interference due to endogenic ammonium was encountered, since the amount of ammonium generated enzymatically is orders of magnitude higher than that naturally present in blood. The activity of the membrane sandwich remained constant during at least one week of continuous daily operation, upon which some deterioration in the measured signal appeared. Since it has been established (19) that this decrease is not due to loss or deactivation of the immobilized enzyme, it is most probably due to clogging of the pores of the hydrophobic gas permeable membrane. Thus another membrane material than that used in this investigation might improve the lifetime of the membranee assembly. Table 1 Assay of Urea in Plasma Samples as Determined at the Copenhagen County Hospital in Herlev (CCHH) and by the Flow Injection Analysis System (FIA) Depicted in Fig. la. All Values are Expressed in mM Urea as a Means of Two Measurements Sample Concentration Determined No. CCHH FIA 1 5.3 5.4 1 o 2 28. 7 26.3 3 18.2 19.7 4 17.1 " 18.5 5 2.4 2.3 6 7.7 8.1 i s 7 20.4 18.6 8 8.8 10.0 9 26.2 24.3 10 18.0 18.3 11 5.0 6.3 2 o 12 33.1 32.1 13 19.4 18.8 14 7.9 8.8 15 9.2 9.3 16 9.0 9. 7 CONCLUSION Urea in whole blood, enzymatically degraded by immobilized urease followed by colorimetric detection of the ammonia generated, can be 5 determined in a FIA system provided that the dispersion coefficient of the injected sample at the point of quantification is as close to unity as possible. With this condition fulfilled, the analytical readout is independent of any ° variations in the pH, the buffering capacity and the hematocrit level of individual samples. Previously it has ben postulated that FIA for principal reasons is unsuitable for measuring of undiluted blood samples (17, 18). Besides s reporting the first successful enzymatic assay of whole blood, this communication analyses the conceptual problems connected with this type of analysis. It has been shown that by exploiting the unique time/space kinetic discrimination of FIA, this approach allows these problems to be resolved. REFERENCES 1. Karube and S. Suzuki, Biochem. Biophys. Res. Com ., 47 (1972) 51. 2. J. Montalvo Jr. and G.G. Guilbault, Anal. Chem., 41 (1969) 1897. 3. D. Kirstein, L. Kirstein and F. Scheller, Biosensors, 1 (1985) 117. 4. K. Yasuda, H. Miyagi, H. Hamada and Y. Takuta, Analyst, 109 (1984) 61. 5. H. McCullock, Clin. Chim. Acta, 19 (1968) 101. 6. N.J. Szuminsky and A.K. Chen, Biotech. Bioeng., 26 (1984) 642. 7. J.P. Joseph, Microchim. Acta, II (1984) 473. 8. J.J. Kulys, V.V. Gurevicieno and V.A. Laurimavicius, Biosensors, 2 (1986) 35. 9. M.A. Arnold and S.A. Glazier, Biotechn. Lett., 6 (1984) 313. 10. T. Okada, I. Karube and S. Suzuki, Eur. J. Appl-. Microbiol. Biotech., 14 (1982) 149. 11. F. Winquist, A. Spetz and I. Lundstrδ , Anal. Chim. Acta, 164 (1984) 127. 12. F. Winquist, A. Spetz and I. Lundstrδm, Anal. Chim. Acta, 164 (1984) 143. 13. F. Winquist, I. Lundstrom and B. Danielsson, Anal. Chem., 58 (1984) 145. 5 14. J. Ruzicka- and E.H. Hansen, Anal. Chim. Acta, 173 (1985) 3. 15. B.A. Petersson, E.H. Hansen and J. Ruzicka, Anal. Lett., 19 (1986) 649. 16. D. Betteridge, W.C. Cheng, E.L. Dagless, P. 0 David, T.P. Goad, D.R. Deans, D.A. Newton and T.B. Pierce, Analyst, 108 (1983) 17. 17. J.J. Harrow and J. Janata, Anal.. Chim. Acta, 174 (1985) 115. 18. J.J. Harrow and J. Janata, Anal. Chim. s Acta, 174 (1985) 123. 19. B.A. Petersson, Ph.D. Thesis (in prep.). 20. A.U. Ramsing, J. Ruzicka and E.H. Hansen, Anal. Chim. Acta, 114 (1980) 165. EXAMPLE 2 ENZYMATIC DETERMINATION OF UREA IN UNDILUTED WHOLE BLOOD BY FLOW INJECTION ANALYSIS USING AN AMMONIUM ION-SELECTIVE ELECTRODE ABSTRACT 5 A flow injection system for the assay of urea in undiluted whole blood is described. Quantification of urea is achieved by means of an ammonium ion-selective electrode furnished with a membrane incorporating immobilized urease o measuring the enzymatically generated ammonium which is directly related to the concentration of the urea. The interference from potassium is reduced by adjusting the potassium concentration in the carrier stream and in the aqueous s calibration solutions to 4.0 mM. However, a complete elimination of this interference is achieved by measuring the actual potassium concentration in the respective sample using an external instrument, thus mathematically 0 correcting for the additional contribution to the signal due to the presence of K . The linear measuring range is 1-40 mM urea with a sample frequency of 40 h~ and a standard deviation of 1% for whole blood samples. The result of the 5 measurement is obtained within 25 sec. from the time of injection. Variations in the hematocrit level of the sample has no effect on the measurement. The method is in excellent agreement with the one used routinely at a local 0 County Hospital. The sensor is stable for more than 30 days. INTRODUCTION Enzyme sensors, or electrodes as they are conventionably called, consists of an electrode covered with a membrane layer containing one or possibly more enzymes. This type of electrodes have proved to be useful in biochemical analysis and have been developed for a variety of substrates. The enzyme catalyses the reaction of the substrate to be measured by generating or o consuming species for which the inner sensor is selective. In most cases, covalently immobilized rather than physically entrapped enzymes are preferred, because covalent immobilization techniques generally offers increased stability 5 of the enzyme activity. Enzyme electrodes have the advantage of being simple, reliable, essentially reagentless and sensitive, besides being highly selective due to the inherent selectivety of the enzymes. o As early as 1969 Guilbault et.al. (1) described the use of an ammonium ion-selective electrode for the determination of urea, based on the enzymatic degradation of urea by urease according to the following reaction scheme CO(NH) 2 +H 2 0+H 3 0 + UreβSe 2NH*+HCO~ However, the nonactin/monactin electrode used provides poor selectivity for.ammonium ions over potassium. An urea electrode of similar composition was used by Yasuda et.al. (2) in conjunction with a FIA system, where the interference from potassium was compensated for by using an additional ammonium ion-selective electrode (without enzyme) for K detection. This system proved to be s linear in the range 3.6-107 mM urea. However, attempts to determine urea in whole blood was impeded due to influence by the hematocrit level of the sample. Although the flow injection analysis technique is o well established for the analysis of homogeneous solutions, the analysis of heterogeneous mixtures, such as whole blood, presents problems relating to the difference in the dispersion pattern of the two types of samples. In a FIA s system it is essential that the dispersion of the sample is constant and reproducible, so that calibration solutions and unknown solutions are treated in the same way. Harrow et.al. (3) has shown that the presence of red blood cells can o lead to errors related to the hydrodynamic effect of the red blood cells on the dispersion. Thus, not being able to control the dispersion of the blood samples with sufficient accuracy, the FIA manifold therefore should, for pratical assays, 5 be operated at a D-value as close as possible to 1, that is, the distance between the injection port and the detection position should be minimized (3, 4). In this paper, a FIA system for routine 0 determination of urea in undiluted whole blood, using an urea electrode consisting of an ammonium ion-selective electrode furnished with a covalently immobilized urease membrane, is described. The interference from potassium was reduced by adjusting the potassium concentration 5 in the carrier stream and in the aqueous calibration solutions to 4.0 mM. However, a complete elimination of this interference was achived by measuring the actual potassium concentration in the respective sample using an io external instrument, thus mathematically correcting for the additional contribution to the signal due to the presence of K . The measurement is not affected by variations in the hematocrit level of the sample. ls THEORY The electrode potential, E, for a monovalent cation X is described by the Nernst equation: E = E n + RT/F • lna . (1) where E is standard electrode potential 0 (temperature dependent), R is the gas constant, T is the absolute temperature (K), F is the Faraday constant and ax is the ion activity - ~ . At 25 C, eqn.(l) can be written as: E = E Q + 59.16 • log a χ (2) s Eqn. (2) is valid at 100% sensitivity. At lower sensitivity, the Nernstian slope of 59.16 mV/decade must be reduced accordingly. The ion activity (a ) is related to the concentration (c ) by the equation a = y (3) where y J x denotes the activity - ~ coefficient which is governed by the total ionic strength, I, of the sample. The ionic strength is defined as: I = 0.5 c i z (4) where c. is concentration of each ionic species present in the sample and z. is the charge of each ionic species present in the sample. The activity coefficient can, up to an ionic _2 strength of 10 M, with fairly good accuracy be calculated according to the Debye-Hϋckel equation: _ι 0 g y χf (Azjfϊ/ d+B i fl ) ) ( 5 ) Here A and B are general constants assuming a value of 0.5115 and 0.3291 respectively at 25 C and a is the specific ion size parameter. Combining egns. (1) and (3) yields: E = Ev_ + RT/F • ln(yJ ) + RT/F • ln(c ) (6) Assuming that the ionic strength does not vary from sample to sample, the y value is constant and eqn. (6) becomes: where Eo = RT/ 'F • ln(yx) and k is the actual slope of the electrode expressed in mV/decade. Thus, if I is constant, the recorded potential E is a linear function of log(c ). If an interfering cation is present in the sample solution, the Nernst equation must be written in the- following extended form (Nicolsky eqn.): E=Eo+RT/ 'F • In(ax+Kxi.aT1' Z i) (8) where a. is the activity of interfering cation, z. is the charge of interfering cation and Kxi. is the selectivity J coefficient, When.Kxi. values are known, it is p e ossible to compensate for interferences by making appropriate calculations. However, it must be remembered that the K value is not a constant. It varies to some degree with the concentration level of the primary ion (analyte) and of the interfering species, sample composition, time and temperature. K values can be determined in the following way: Firstly, the electrode potential is measured against a suitable reference electrode in a solution where the activity of the primary ion (analyte) is a (E. ) . Secondly, the electrode potential is measured in a solution where the activity of the primary ion (analyte) is identical to that of the first solution and the activity of the interfering ions is a.(E 2 ). When eqn. (1) and eqn. (8) are combined and solved with respect to K . , the following expression is obtained: where ΔE=E 2 -E, , Assuming that I does not vary from sample to sample, eqn. (9) becomes: Thus, a relationship ^ between Kxi. and cx can be determined and used to compensate for interferences if the concentration/activity of the interfering ion is known. EXPERIMENTAL Apparatus All measurements and experiments were executed in the single line FIA manifold depicted in Fig. 3, comprising a flow through reference electrode (El -Radiometer A/S, K1702), a rotary injection valve, and an urea electrode (E2) . The sample solution (45 l) is injected into the carrier stream, the distance between the injection port and the 5 μl flowcell (FC2) being made as short as possible in order to ensure minimum dispersion (D value) of the sample zone. The urea electrode was made by placing an immobilized urease membrane (cf. below) on top of a small ammonium ion-selective electrode consisting of a actin/monactin PVC membrane. The diameter of the active surface was 1.5 mm. The enzyme electrode was connected to a Radiometer A/S PHM 64 pH meter and the potentials were recorded on a Servograph REC 61 recorder furnished with a REA 112 high-sensitivity interface (Radiometer A/S), the peak height at peak maximum serving as the analytical readout. The carrier stream was propelled by means of a Radiometer A/S ABU80 syringe pump (P2) at 0.125 ml/min.. The concentration of potassium in actual whole blood samples was measured with a KNa instrument from Radiometer A/S. All connections were made with microline tubings (o,5 mm i.d. ) . Reagents The carrier stream solution consisted of an isotonic HEPES buffer comprisning 4 mM KCl, 86 mM NaCl, 60 mM Na-HEPES and 0.55 g NaEDTA/L. All chemicals were of analytical-reagent grade. The aqueous urea standards in the. range 1-40 mM used for calibrating the system were prepared by dissolving appropriate amounts of urea in the isotonic HEPES buffer. Human blood were obtained from the clinical chemistry department of a local hospital (Copenhagen County Hospital in Herlev) . Preparation of Urea Electrode 40 mg of urease from "jack beans" (E.C. 3.5.1.5, Sigma, 70 U/mg) was dissolved in 0.5 ml phosphate buffer (10 mM, pH 7,2) and 25 μl glutaraldehyde (25% aqueous solution from Sigma) was added. The solution became viscous within a few min., whereupon 20 μl of this mixture was placed on a piece (6 mm in diameter) of polyester membrane . (Polymon, PES-50/36, thickness 41 im) . The immobilized urease membrane was fixed on the ammonium ion-selective electrode by means of an o-ring. When not in use the membrane was stored in a refrigerator (4 C) where it is stable for several months without any detectable deterioration. RESULTS AND DISCUSSION All measurements and experiments were exceuted in the sigle line FIA manifold depicted in Fig. 3. Note that the flow through reference electrode is placed immediately before the injection port, so that the risk of clogging of the porous pin, due to absorption of proteines from the blood samples, is avoided. Hemolysis tends to occur under hypotonic conditions. Thus, an isotonic buffer (1-0.15 M) solution should be employed as carrier stream to prevent the whole blood from hemolysing. A phosphate buffer is often used in conjunction with enzyme analysis because it is compatible with many enzymes. However, phosphate has the ability to form ammonium complexes, thus reducing the analytical signal due to masking effect. In order to minimize the liquid junction potential, i.e., the unavoidable potential difference between the sample solution and the carrier solution existing at the liquid-liquid phase boundary, a biological HEPES buffer was chosen as carrier stream. Since the normal physiological concentration of potassium is very close to 4.0 mM the concentration of potassium in the carrier stream was adjusted to this concentration, thus reducing the interference from this analyte and stabilizing the background potential. It was investigated whether a constant exposure of the ammonium ion-selective electrode (ASE) to potassium had any adverse influence on the selectivity coefficient as a function of time. However, the selectivity coefficient remained constant during 1 month of operation. In order to prevent any inhibition of the urease enzyme due to presence of different metal ions, 0.55 g NaEDTA/L was added to the carrier solution. At a total ionic strength around 0.15 M the "naked" ammonium ion-selective electrode (ASE) exhibits a linear calibration curve in the range 10 -3- 100 M ammonium. Since the physiological range of urea is 1-40 mM it is obvious that the enzymatic conversion of urea to ammonium has to be greater than 50% (1/2 mol of urea is converted into 1 mol of ammonium) to obtain an ammonium concentration within the linear range of the ASE. The possibility of producing a membrane with an urease enzyme activity capable of converting more than 50% of the urea in the sample was first evaluated by placing a cuprophane membrane with physically entrapped urease on the ASE. At a flow rate of 0.125 ml/min the linear range of this electrode was 1-40 mM urea, the electrode initially exhibiting a slope of 41 mv/decade. However, the slope of the enzyme electrode decreased with time (cf. Fig. 2) due to poor stability of the enzyme membrane. When the slope of the enzyme electrode decreases the selectivity coefficient increases (cf. eqn. (10)). Since the selectivity coefficient is a function of the urea o concentration several standard solutions are needed to determine this relationship. Thus, aiming at manufacturing a commercially viable urea electrode it is essential that the slope of the sensor remains relatively constant during approximately 1 month of operation in order to ensure that the relationship between the concentration of the analyte and K. , which for practical reasons has to be preprogrammed into the instrument, does not change. Encouraged by the preliminary experiments, a membrane containing covalently immobilized urease (cf. above) was made and tested in the FIA system. The linear range of the sensor was 1-40 mM urea with a sample frequency ' of 40 h~ and a standard deviation of 1%. The enzyme membrane was found to be stable for more than 1 month due to the increased stability of the covalently immobilized enzymes. The variations of K 1. from electrode to electrode was evaluated with and without the incorporation of an enzyme membrane using 5 different electrodes (ASE.-ASAg). The K. value was only determined in the range 1-10 M urea and 1-10 mM ammonium. At higher concentration levels the interference from potasiu is of less importance. As expected, the "naked" ASE can be made very reproducible (cf. Table 2) owing to the uniformity of the PVC ion-selective membranes. TABLE 2 Determination of K. as a function of the 1 ammonium concentration using 5 different electrodes without enzyme membranes. ASE, ASE 2 ASE 3 ASE 4 ASE 5 K- 0.14 0.14 0.14 0.14 0.14 K 1.0 0.21 0.21 0.20 0.21 0.21 K indicates the K value at an ammonium x concentration of x mM 5 different urease membranes were prepared as reproducibly as. possible and the relationship between the urea concentration and the K. value was established for the corresponding urea electrodes (cf. Table 3). The similarity of K. from one enzyme electrode to another is quite sufficient to allow the instrument to have a preprogrammed function for a mathematical correction of the interference from potassium. The actual electrode slope was in all cases found to be within the range 42.610.9 V/decade. Whole blod is an inhomogeneous liquid and its dispersion pattern in a FIA system is completely different to that of a homogeneous liquid due to the blood cells ability to act as very efficient stirring elements. Minimum hematocrit effect is achieved, when operating at a D-value as close as possible to 1, that is, the distance between the injection port and the detection position should be minimized (3, 4). In the FIA system depicted i o in Fig. 3 this distance amounted to a volume of 12 μl . TABLE 3 Determination of K. as a function of the urea concentration using 5 different electrodes with ! 5 enzyme membranes. 0 Kx indicates the K value at an urea concentration of x mM A typical calibration graph is obtained with a series of aqueous urea standards. At urea 5 concentrations higher than 40 mM the calibration graph is no longer lineari The curvature of the calibration plot at concentrations above 40 M urea might be due to several factors. In order to obtain a linear relation between the recorded 0 potential and the log (urea) (cf. eqn. (7)) the following three conditions must be fulfilled: (a) The kinetics of the urease enzyme has to follow that of pseudo-first order reaction conditions, i.e. the amount of urease must be in excess and CureaD « K-, s (Michaelis-Menten constant). (b) The change of pH within the enzyme membrane due to formation of ammonium must be limited in order to maintain the optimum pH value for the enzyme reaction during the i measurement. (c) The change of ionic strenght within the enzyme membrane due to formation of ammonium ions must be limited, to fulfil the assumption of constant ionic strength is (cf. the discussion in the theory section). Further optimization of the system with respect to increasing the linear range was not attempted since a linear range of 1-40 mM urea is sufficient to cover the physiological range. 0 However, as described by Petersson (5), an increase of the linear range for an enzyme electrode is readily achieved in a FIA system. In order to examine the influence of the hematocrit level of the sample donor but with 5 different amounts of red blood cells were prepared and injected into the FIA system. The results showed (Table 4) that the influence of the hematocrit value on the analytical readout was minimal. TABLE 4 The influence of the hematocrit level of the sample on the measurement. Hematocrit % mM urea The coefficient of variation for 10 measurements l o of a blood sample was 1.06%. A closer inspection of the response curves obtained with blood samples and aqueous standard solutions, respectively, revealed that the dispersion of the two types of samples were not 1 5 identical, i.e., the time span between the injection of the sample and the beginning of the response curve was different. It is well known that the dispersion process is a function of the viscosity of the sample (5, 6). Generally 0 speaking, with increasing viscosity of the injected sample, the zone broadening and ^ hence the dispersion decreases. Thus, the viscosity was increased to a level comparable with that of a blood sample by addition of 50 mg albumin/ml to 5 a 10 mM aqueous standard solution. The dispersion of this sample was found to be identical to that of a whole-blood sample. However, the analytical readout measured as the peak height at peak maximum was not influenced by the viscosity. Thus, the dispersion of the fluid element corresponding to peak maximum is not 5 influenced by changes in the viscosity of the sample and aqueous standard solutions can therefore be used for calibration .of the system. Next, a series of 41 whole blood samples, supplied by the Copenhagen County Hospital in 1.Q Herlev (CCHH), were assayed in order to compare the FIA method with that used routinely at CCCHH where urea exclusively is analyzed in plasma. The results are reproduced in Table 5, the correlation between CCHH and FIA = 0.95 CCHH-0.11 is , with a correlation coefficient of 0.999. The average standard deviation (n=2) was 0.86%. The enzyme electrode was found fully compatible with whole blood samples, i.e., the slope of the electrode showed no change during the injection 2o of more than 150 whole blood samples into the FIA system. The concentration of potassium in the actual whole blood samples were 1 measured using an external static apparatus and mathematical 25 corrections for this interference were made. When the concentration of potassium is within the normal physiological range only small corrections (<__3%) are needed, due to the background K-level maintained in the carrier stream and the aqueous 3 standard solutions. However, since the physiological range is "2-9 M K , corrections TABLE 5 Assay of Urea in Whole Blood Samples as Determined at the Copenhagen County Hospital in Herlev (CCHH) and by the Flow Injection Analysis System (FIA) Depicted in Fig. 3. All Values are Expressed in mM Urea as a Means of Two Measurements. The correlation coefficient is 0.999. up to - 15% can occur. Even though the actual determinations of the potassium have been made in a static system, there is no reason, whatsoever, that a FIA system similar to that used for the derermination of urea, should not be used in conjunction with a potassium ion-selective electrode, thus making a parallel determination of these two analytes. Since urea and potassium often are determined together a simultaneous determination of these two parameters would, from a clinical point of view, be desirable. A STAT system for quantification of urea and K in whole blood samples based on the FIA technique is a very realistic possibility. CONCLUSION In the present work a FIA system for practical determination of urea in undiluted whole blood samples using an ammonium ion-selective electrode coupled with an urease enzyme membrane has been applied. The results obtained with this system were in excellent agreement with the method routinely used at the Copenhagen County Hospital in Herlev (CCHH) where urea exclusively is analysed in plasma (FIA = 0.95 CCHH-0.11 mM, with a correlation coefficient of 0.999). The linear range was 1-40 M with a sample frequency of 40 h and a standard deviation of 1% for whole blood samples. The result of the measurement is obtained within 25 sec. from the time of injection. Variations in the hematocrit level of the sample has no or minimum effect on the analytical readout. The interference from potassium was reduced ( <±3% within the normal physiological range) by adjusting the potassium concentration in the carrier stream and in the aqueous calibration solutions to 4.0 mM. However, a complete elimination of this interference was achieved by measuring the actual potassium concentration in the respective sample using an external instrument, thus mathematically correcting for the additional contribution to the signal due to the presence of K . A FIA system for parallel determination of urea and potassium is a realistic possibility and since there exists a clinical requirement for an instrument which facilitates a simultaneous determination of K and urea, a STAT-FIA approach should be very attractive. The present work is to the author's knowledge the first of its kind with respect to the use of the FIA technique for practical determination of species in whole blood. However, there is no doubt that commercial FIA based instruments for clinical applications will be purchasable within the next 10 years. A commercial FIA instrument for the assay of plasma glucose has, in fact, already ben marketed by the Japanese company Shi adzu. REFERENCES 1. J. Montalvo, Jr., and G.G. Guilbault, Anal Chem., 41 (1969) 1897 2. K. Yasadu, H Miyagi, Y. Hamada and Y. Takata, Analyst, 109 (1984) 61 3. J. Harrow and J. Janata, Anal. Chim. Acta, 174 (1985) 115 4. B.A. Petersson, H. Andersen and E.H. Hansen, Anal. Lett., 20 (1987) (in press) 5. B.A. Petersson, Ph.D. Thesis 6. D. Betteridge, W.C. Cheng, E.L. Dagless, P. David, T.B. Goad, D.R. Deans, D.A. Newton and T.B. Pierce, Analyst, 108 (1983) 17 EXAMPLE 3 AMPEROMETRIC ASSAY OF GLUCOSE AND LACTIC ACID BY FLOW INJECTION ANALYSIS ABSTRACT A computer controlled flow ' injection system for the assay of D-glucose and L-lactic acid in undiluted plasma is described. Quantification of glucose/lactate is achieved by coupling of an immobilized glucose oxidase/lactate oxidase enzyme membrane with an amperometric electrode, thus measuring the enzymatically generated hydrogen peroxide, which is directly related to the concentration of glucose/lactate. The amperometric electrodee consists of a platinum anode, to which a potential of +600 mV is applied, and a Ag/AgCl cathode. The diameter of the exposed electrode surface is 2.5 mm. The linear range is 0-40 mM and 0-10 mM for glucose and lactic acid respectively. The sample ffrreeqquueennccyy iiss 66C0 h with a standard deviation of less than 1.5%. Potential application of the FIA system as a STAT glucose analyzer for whole blood measurements is discussed. INTRODUCTION The detection, identification and quantification of glucose in blood has played a vital role in the diagnosis and management of patients suffering from disorders of carbohydrate metabolism and is one of the most frequently performed determinations in clinical chemistry. Analysis of blood glucose are used to detect hyper- and hypo-glycaemia, both of which can result from a variety of different disorders, many of them of endocrine origin. The most common cause of hyperglycaemia is diabetes mellitus, and elderly patients may present in hyperosmolar non-ketotic coma with blood glucose levels as high as 60 mmol/L. Apart from diagnosis, serial blood glucose levels are required to monitor diabetes on treatment, and to detect and ultimately to prevent gestational hyperglycaemia which although not necessarily harmful to the mother, may be damaging to the foetus (1) . In patients with acute perfusion failure and for critically ill and injured patients in general, the arterial lactate concentration is now recognized as a very good indicator of severity and prognosis (2, 3). The increases in blood lactate are due to a critical reduction in oxygen delivery such that aerobic metabolism through the Krebs cycle cannot be sustained. The severity of the defect in oxygen delivery is proportional to the increase in lactate generated through the emergency pathway of glucose metabolism from pyruvate. Numerous reports concerning amperometric determination of glucose using glucose oxidase (GOD) to catalyze the reaction: /3-D-glucose + 0 2 +H 2 0 > D-gluconic acid + H 2 0 2 thus measuring the 0 2 consumption by a Clark electrode or the formation of hydrogen peroxide by a reversed Clark electrode, has been published (4-16) . Similar electrodes for quantitative determination of L-lactate based on the enzymatic degradation of L-lactate by lactate oxidase (LOD) according to the following reaction: L-lactate + 0 2 +H 2 0 L0D CH 3 COCOOH +H 2 0 2 * has been reported (17-20). The linear ranges of enzymatic methods are normally limited by the Michaelis constant (FC of the enzyme, and sample dilution is usually necessary for clinical assays. By including an outer membrane with low permeability to restrict diffusion of glucose to the underlying glucose oxidase layer a linear range of up to 50 mmol/L glucose is achieved (16). Another and more recent approach has been to use ferrocene as an electron mediator between glucose oxidase and the working electrode, thus expanding the linearity up to 30 mmol/L glucose (21). However, it is still a fact that, even though ampero etric glucose electrodes have been discussed for almost 20 years, a commercial glucose electrode based system, which can measure on whole blood, is yet to come. In the present work a computer controlled FIA system for enzymatic determination of glucose or lactate, based on the ultimate generation and detection of hydrogen peroxide, using GOD or LOD membranes respectively, coupled with an amperometric electrode, are described. The inherent lack of selectivity of the amperometric H 2 0« electrode is compensated for by exploiting the specificity of a cellulose acetate membrane, thus suppressing the interference of substances in blood samples, e.g. glutathione, ascorbic acid, uric acid and paracetamol. The measurements are not affected by variations in the pO_ level of the sample solution. EXPERIMENTAL Apparatus Samples (25 μl ) were injected into the single line FIA system depicted in Fig. 4 by means of a 5 rotary valve operated by a step motor. The distance between,the injection port and the 10 JJ.1 flow cell (FC) was made as short as possible in order to ensure minimum dispersion of the sample zone. o A special designed electrode (GE/LE) comprising a central 0.4 mm diameter platinum working electrode with an outer 2.5 mm-diameter silver ring as the reference electrode was polarized at + 600 mV and used for hydrogen peroxide detection. 5 The electrode system was connected to a home-built potentiostat and the current intensity as a function of time was recorded by a BBC computer from Acorn DFS utilizing the built-in analogue to digital facility. The carrier stream 0 was propelled by means of a Radiometer A/S ABU80 syringe pump (PI) at 1.0 ml/min. The BBC computer facilitated the data aquisition and the operation of the sampling pump, the syringe pump and the injection valve. s All connections were made with microline tubing (0.5 mm i.d. ) . REAGENTS The isotonic phosphate buffer used as carrier stream was prepared by dissolving 1.12 g of NaEDTA, 1.84 g NaBenzoate, 3.04 g NaH 2 P0 4 , 13.80 g Na 2 HP0 4 and 5.44 g NaCl in 2 1 of distilled water. Aqueous glucose and lactate standards were prepared from stock solutions containing 100 mM glucose and 100 mM lactate respectively. For the amperometric enzyme electrodes, high rejective glucose or high rejective lactate membranes (3.0 mm diameter) purchased from Yellow Springs Instrument Co. was placed on the electrode tip and fixed with an outer polycarbonate membrane (Nucleopore, 5 ^μm thick, 0.08 urn pore size) held with an o-ring. RESULTS AND DISCUSSION All experiments and measurements were executed in the FIA system depicted in Fig. 4. A BBC computer was used to facilitate the following operations: 1) a one point calibration (mM) 2) data aquisition 3) a graphic display of the response curve 4) control of the sampling and syringe pump 5) presentation of data, i.e. the concentration in mM, the time of peak maximum and the sample no. 6) manual injection of sample 5 7) a complete storage of data, which of cause can be recalled. The enzymatic degradation of a substrate (S) to a product (P) can in many cases be described by simple Michaelis-Menten kinetics o kl S + E = ES ~-→ P + E where the substrate is bound reversibly to the enzyme (E) to form a substrate-enzyme complex (ES), which decomposes irreversibly to the product and the enzyme with the rate coefficient 5 kl. The rate of formation of product is where CE Q 0 = (X) + CES3 and K M is the Michaelis constant with the dimension of concentration. If S « K.,, eqn. (1) simplifies 0 to a first order expression: d(P)/dt = (klCE^/K j ,) - (S) (12) Eqn. (11) is not valid for enzyme systems which require additional cosubstrate(s) or coenzyme(s). Nevertheless eqn. (11) has proved 5 to be a very useful approximation as long as the cosubstrate(s) or coenzyme(s) are in excess. In order to obtain a linear relationship between the glucose or the lactate concentration and the recorded signal the two following conditions must be fulfilled: (a) The kinetics of the glucose oxidase or the lactate oxidase enzyme has to follow that of a pseudo-first order reaction conditions. (b) Oxygen must be in excess. The physiological range of glucose and lactate is larger than K M for each of the two enzymes. Thus, in a static system, first order kinetics can only be achieved by including an outer membrane with low permeability to restrict diffusion of the analyte to the underlying enzyme layer. However, in a FIA system, restricted diffusion of analyte to the enzyme can be controlled by the time of exposure of the sample to the electrode, the dependent variable being the flow rate. This is illustrated in Fig. 5, where the linear range of the glucose electrode is expanded from 20 to 40 mM glucose by increasing the flow rate from 0.5 to 1.0 ml/min. The influence of the oxygen level on the measured glucose value was investigated. A series of aqueous glucose standards containing 10, 20 and 30 mM glucose was prepared and the pO„ level was adjusted in the range 1-140 mm Hg by means of a tonometer. The samples were injected into the FIA system and the glucose concentration as a function of p0 2 was monitored (cf. Table 6). It might appear that there is a discrepancy between the statement that there is a minimum dispersion of the sample zone, while at the same time it is obvious that the oxygen level of the sample has no influence on the measured glucose concentration. However, this is due to yet another feature of the FIA technique, that is, the perfectly controlled spatial and kinetic discrimination prevailing in the system (22). TABLE 6 The glucose concentration as a function of p0 2 - p0 2 (mm Hg) 10 M glucose 20 mM glucose 30 M glucose At the start of the measurement, glucose and hydrogen peroxide are not present in the enzyme membrane which is equilibrated with oxygen from the carrier solution. When the injected sample reaches the enzyme electrode, glucose diffuses into- the enzyme membrane and the degradation of glucose takes place, while the undiluted sample zone is swept away by the carrier stream. In hydrophilic gel membranes, oxygen diffuses between three and four times faster than glucose (23,24). Since the time of exposure of the sample to the electrode is very short the oxygen from the carrier stream is transferred to the enzyme membrane so fast, that the oxygen concentration in the membrane is always close to the oxygen concentration of the carrier solution. Thus, it is possible to disregard the p0 2 level of the sample as a limiting factor with respect to the enzymatic reaction. Potential interferences from ascorbic acid, glutathione and paracetamol were investigated but no effects on the glucose response was found due to the selectivity of the cellulose acetate membrane included in the YSI membrane sandwich. Finally, the extent of variation in the dispersion as a function of the viscosity was evaluated by measuring the glucose concentration in aqueous samples containing 0-160 mg albumin/ml and a fixed amount of glucose. The glucose measurements - and hence the dispersion process - were not affected by the addition of albumin. The dispersion, if any, is influenced by the viscosity of the sample, thus leading to a change in the analytical readout. Therefore it may be concluded that there is no or minimum dispersion in the FIA system depicted in Fig. 4. A typical calibration run is obtained with a series of aqueous glucose standards. It is worth while to notice that the result of the analysis is completed within 5 sec. from the time of injection. The linear range is 0-40 mM glucose with a sample frequency of 60 h " and a standard deviation of less than 1.5%. Next, a series of 34 blood plasma samples, supplied by the Copenhagen County Hospital in Herlev (CCHH), were assayed in order to compare the FIA method with that used routinely at CCHH. The results are reproduced in Table 7, the correlation between CCHH and FIA being CCHH = 0.98 FIA + 0.07 mM, with a correlation coefficient of 0.985. It is feasible to increase the sampling frequency to approximately 120 h~ by applying a normal glucose oxidase membrane (YSI). However, it is possible for paracetamol to penetrate the thin (1.2 jum) cellulose acetate layer and therefore give rise to interference. Thus, it was investigated whether the potential could be lowered in order to prevent oxidation of paracetamol at the platinum working electrode. The potential was varied in the range +200 to +600 mV and it turned out that an oxidation of paracetamol can be prevented at a potential of +300 mV (cf. Fig. 7) . Normally a potential of 600-650 mV is chosen because the current is not affected by small changes of the potential in this range. However, a potential of +300 V can be used in conjunction with a three electrode system. The high rejective glucose oxidase membrane was then replaced with a high rejection lactate oxidase membrane, and a series of 11 blood plasma samples, supplied by the Copenhagen County 5 Hospital in Herlev (CCHH), were assayed in order to compare the FIA method with the YSI lactate analyzer. The results are reproduced in Table 8, the correlation between YSI and FIA being YSI = 1.04 FIA-0.02 mM, with a correlation coefficient 0 of 0.967. TABLE 7 Assay of Glucose in Plasma Samples as Determined at the Copenhagen County Hospital -in Herlev (CCHH) and by the Flow Injection Analysis System (FIA) Depicted in Fig. 4 . All Values are Expressed in mM 5 Glucose as a Means of Two Measurements o 5 o Assays on undiluted whole blood were not possible due to a very bad reproducibility (±10%) caused b y the stirring effect of the blood cells. However, it has been shown by Petersson (25) that b y proper design of the FIA system, it is accomplishable to measure urea in whole blood samples - the standard deviation being 1.0%. It may very likely be possible to measure glucose/lactate on whole blood samples if the geometry of the amperometric glucose/lactate electrode and the flow cell is optimized. TABLE 8 Assay of Lactate in Plasma Samples as Determined by the YSi Lactate Analyzer and by the Flow Injection Analysis System (FIA) Depicted in Fig. 4. All Values are Expressed in mM Lactate as a Means of Two Measurements o o CONCLUSION In the present work a FIA system has been applied for detection of glucose and lactate in plasma using an amperometric electrode with detection of enzymatically generated hydrogen peroxide. The results were in excellent agreement with other methods, i.e. CCHH = 0.98 FIA + 0.07 mM glucose and YSI = 1.04 FIA-0.02 mM lactate. The analysis is completed within 5 sec. and there is no interference from other species. The linear range is 0-40 mM glucose and 0-10 mM lactate with aa ssaammppllee ffrreeqquueennccyy ooff 6600 hh~~ and a standard deviation of less than 1.5%. The very fast response time makes this system attractive as a STAT glucose analyser if accommodation of whole blood samples were possible. Further optimization of the geometry of the system is necessary in order to reduce the error related to the effect of red blood cells on the dispersion. REFERENCES 1. J.M. Burrin and C.P. Price, Ann.Clin.Biochem. , 22 (1985) 327 2. L.D. Cady, M.H. Weil and A.A. Afifi, Crit. Care Med., 1 (1973) 75 3. G . Broder and M.H. Weil, Science, 143 (1964) 1457 4. L.B. Wingard, Jr., D. Ellis, S.J. Yao, J.G. Schiller, C.C. Liu, S.K. Wolfson, Jr. and L. Drash, Journal of Solid-Phase Biochemistry, Vol. 4, No. 4 (1979) 253 5. M. Kessler, J. Hδper, H.J. Volkholz, D. Sailer and L. 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Claims: CLAIMS 1. A method of determining the concentration of a primary analyte in each of a sequence of fluid samples by means of a sensor responsive to 5 a secondary analyte derived from the primary analyte in one or a plurality of chemical reaction steps, c h a r a c t e r i z e d by providing at the test location a stationary fluid element between o the actual fluid sample under test and the sensor; providing at the fluid element a chemical system reactive with the primary analyte with a resulting formation of the secondary analyte; contacting the fluid element in a controlled time s with the actual fluid sample under test; time controlled exposure of the fluid element prior to and subsequent to the contact with the actual fluid sample under test to a conditioning medium of a well defined chemical composition, o thereby maintaining the fluid element as a matrix essentially independent of the chemical composition of the actual fluid sample under test, in which matrix the conversion from primary analyte to secondary analyte takes place. 5 2. A method of determining the concentration of an analyte in each of a sequence of fluid samples by means of an analyte sensor whereby a sample with a fluid sample analyte concentration greater than the upper concentration limit for the sensor linearity range irregularly may occur, c h a r a c t e r i z e d by providing at the test location a stationary fluid element between the actual liquid sample under test and the analyte sensor; controlling the time of exposure of the actual fluid sample under test to the fluid element to 1 o be that short that a fluid sample of the type showing an analyte concentration between the upper concentration limit for the sensor linearity range and the maximum expectable analyte concentration will provide at the time of is test an analyte concentration in the fluid element lower than the upper concentration limit for the sensor linearity range. 3. A method of determining the concentration of an analyte in each of a sequence of fluid zo samples by means of an analyte sensor whereby a sample comprising an interfering chemical species with lower diffusion characteristics than the analyte irregularly may occur, c ha r a c t e r i z e d by providing at the s test location a stationary fluid element between the actual fluid sample under test and the sensor; controlling the time of exposure of the fluid sample under test to the fluid element to be that long that at the time of test an analyte concentration measureable with the analyte sensor is established in the fluid element and at the same time that short that the concentration of the interfering species established in the fluid element be low and without significant effect on the analyte sensor response. 4. A method according to any of claims 1-3, c h a r a c t e r i z e d in that the individual samples of the same sequence of fluid samples comprise samples with differing matrices such as samples of biological or physiological origin, in particular whole blood. 5. A method according to claim 1, c h a r a c t e r i z e d in that the primary analyte is urea, the secondary analyte is ammonia, the reactive chemical system is the enzyme urease and the conditioning medium is a pH buffer fluid of an alkaline pH. 6. A method according to claim 1, c h a r a c t e r i z e d in that the conditioning medium is a fluid of well defined oxygen partial pressure, e.g. ambient oxygen partial pressure, and that oxygen is consumed in the conversion of the primary analyte to the secondary analyte. 7. A method according to claim 2, c h a r a c t e r i z e d in that the analyte is a metabolite, e.g. glucose or lactate. 8. A method acording to claim 3, c h a r a c t e r i z e d in that the analyte is s glucose and the interfering species is paracetamol. 9. An analyzer for performing any of the methods according to the preceding claims. 10. An analyzer according to claim 9, c h a r a c t e r i z e d by being a flow injection analyzer. 11. An analyzer according to any of claims 9-10, c h a r a c t e r i z e d by being operated in a non dispersion mode, i.e. that with respect to a flow injection analyzer the value of the dispersion coefficient, D, is preferably lower than 1.1 and in particular lower than 1.05. 12. An analyzer according to any of claims 9-11, c h a r a c t e r i z e d in that the support means for the stationary fluid element has such dimensions that the distance between the sample under test and the sensor is not more than 1 cm, preferably not more than 1 mm, more preferably not more than 500 μm, for example from 5 μm to 200 μm.

Inventors: PETERSSON BO ARNE, PETERSSON, BO, ARNE
Applicants: RADIOMETER AS, RADIOMETER A/S
Application year: 1988
Main classification: G01N35
IPC codes: C12Q1