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Urea amperometric biosensors based on a multifunctional bipolymeric layer: Comparing enzyme immobilization methods
Sensors and Actuators B 137 (2009) 476–482
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Sensors and Actuators B: Chemical journal homepage: www.elsevier.com/locate/snb
Urea amperometric biosensors based on a multifunctional bipolymeric layer: Comparing enzyme immobilization methods Mariana P. Massafera, Susana I. Córdoba de Torresi ∗ Instituto de Química, Universidade de São Paulo, C.P. 26077, 05513-970 São Paulo (SP), Brazil
a r t i c l e
i n f o
Article history: Received 29 July 2008 Received in revised form 29 January 2009 Accepted 5 February 2009 Available online 20 February 2009 Keywords: Biosensors Urease Poly(pyrrole) Poly(5-amino-1-naphthol) Urea Ammonia
a b s t r a c t This paper outlines the results obtained with biosensors designed for urea amperometric detection. The incorporation of urease into a bipolymeric substrate consisting of poly(pyrrole) and poly(5-amino-1naphthol) was performed through four different approaches: direct adsorption, entrapment in cellulose acetate layer, cross-linking with glutaraldehyde, and also covalent attachment to the polymeric matrix. Poly(pyrrole) acts as amperometric transducer in these biosensors, while poly(5-amino-1-naphthol) drastically reduces the interference signal of agents such as ascorbic and uric acids. The biosensors containing urease covalently attached to the substrate provided interesting results in terms of sensitivity towards urea (0.50 A cm−2 mmol−1 L), lifetime (20 days) and short response times, due to the enzyme immobilization method used. All biosensors analyzed showed also a wide linear concentration range (up to 100 mmol L−1 ) and low detection limits (0.22–0.58 mmol L−1 ). © 2009 Elsevier B.V. All rights reserved.
1. Introduction In human organism, as urea is a final product of protein metabolism, its accumulation above critical levels must be carefully determined. So, urea monitoring is of great interest in clinical and biomedical analysis because this biomolecule enables the diagnosis of some diseases such as kidney and liver malfunction, and also make possible to predict the nature and causes of diabetes [1–3]. Urea is also related to other processes such as soil fertilization, in which it acts as a nitrogen source for plants. For the purpose of urea sensing, various procedures are applied and interesting results have been obtained with enzyme-based sensors. The incorporation of a biological moiety (enzymes, cells, antigens, DNA) to a conventional physico-chemical transducer characterizes a biosensor, and the presence of the biomolecule provides a high specificity to biosensors. The enzyme used in urea biosensors, urease, catalyses the hydrolysis of urea producing ammonium and bicarbonate ions, causing a pH increase in the aqueous reaction medium, as shown in the following reaction: urease
NH2 CONH2 + 3H2 O −→ 2NH4 + + HCO3 − + OH−
(1)
Considering this reaction, the determination of urea can be indirectly done by using specific transducers that detect ammo-
∗ Corresponding author. Tel.: +55 11 30912350; fax: +55 11 37855579. E-mail address: [email protected] (S.I.C.d. Torresi). 0925-4005/$ – see front matter © 2009 Elsevier B.V. All rights reserved. doi:10.1016/j.snb.2009.02.013
nium ions [4], ammonia gas [5], carbon dioxide [6], or pH changes [2–3]. Alternatively, techniques as calorimetric [7] and fluorimetric analysis [8–9] are also available, but one disadvantage of these methods is that in most cases, the sample must be diluted and pretreated before analysis. However, biosensors containing electrochemical transducers are the most used and in these cases urea is detected via either potentiometric [10–12] or amperometric methods [1,3,13–14]. This last one has attracted a lot of attention due to both its effectiveness and simplicity of preparation. Regarding electrochemical biosensors, conducting polymers can be employed as very suitable transducers. An outstanding characteristic of these polymers is their ability to reversibly alternate between the conducting and insulating electronic states. This alternation occurs because conducting polymers present in their structure sequences of conjugated -bonds, changing from insulating to conductors by an oxidation or reduction process, in which incorporation of ions can take place (doping). Many other properties of these special polymers support their application in biosensors, but mainly their ease of preparation (direct deposition of a polymeric film onto an electrode can be performed by electrochemical oxidation of monomers) and other controlled parameters (film thickness, kind of dopant). Among conducting polymers, poly(pyrrole) (PPy) plays an important role because it presents extremely favorable characteristics, as for example its easy synthesis, good electrochemical reversibility, excellent stability, besides presenting a good electroactivity in a large pH range [15–19]. Another important feature of conducting polymers is that some of them can act as selective layers, reducing the amperometric
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signal of interferents as ascorbic and uric acids, species which are often present in clinical samples. This is the case of poly(5-amino-1naphthol) (poly(5-NH2 -1-NAP), Scheme 1), a polymer that was used in addition to PPy in the development of ammonia amperometric sensors [20–21]. Poly(5-NH2 -1-NAP) presents a “polyaniline-like” structure with free –OH groups [22–23] that can undergo a chemical reaction to provide a covalent linkage between the polymeric chain and an enzyme, for example urease, as it will be shown in this work. Covalent attachment of the biomolecule to the substrate is one of the most elegant immobilization methods available, but others as adsorption, entrapment and cross-linking are often used [24]. As will be discussed later, the covalent attachment of the biomolecule to the transducer presents the advantage of enhancing biosensors stability, once one of the major problems of biosensors preparation is the leaching of the biomolecule out of the substrate. The purpose of this work is to provide a systematic comparison among urea amperometric biosensors prepared through those four immobilization methods explained above, using as biological recognition element the enzyme urease. The biosensors consist of a bipolymeric substrate of PPy and poly(5-NH2 -1-NAP), the first acting as electrochemical transducer for ammonia detection (as shown in Fig. 1a), substance that is produced during enzymatic hydrolysis of urea in alkaline medium (reaction (1)). The second plays a double role once it is necessary to covalently attach urease to the substrate and also to reduce interferents signal. 2. Experimental 2.1. Reagents Pyrrole (Aldrich, 99%) was purified by distillation and kept in refrigerator under N2 atmosphere, in the dark. Dodecilbenzenesulfonate, sodium salt (DBSA), 5-amino-1-naphtol (5-NH2 -1-NAP) 97%, cellulose acetate, cyanuric chloride 99% (Aldrich), NaOH, HCl, H3 BO3 (Synth, analytical grade), glutaraldehyde 25% v/v (Merck), NH4 Cl (Mallinckrodt, analytical grade), l-ascorbic and uric acids, bovine serum albumin (BSA) 200 mg mL−1 and Urease (EC 3.5.1.5, type III from Jack Beans, 29,500 units g−1 ) (Sigma) were used as received. All the solutions were prepared using water purified by an Elga UHQ system. 2.2. Instrumental Monomers electropolymerization and chronoamperometric tests (both ammonia and urea detection) were performed in a three electrode conventional cell consisting of a working electrode (Pt/PPy for ammonia detection, and Pt/PPy/poly(5-NH2 -1-NAP) modified with urease in case of urea determination), a Pt foil as counter-electrode and Ag/AgCl/KClsat as the reference electrode. The electrochemical data were obtained via an Autolab equipment model PGSTAT 30 (Eco Chemie), interfaced with a computer for data acquisition.
Fig. 1. (a) Urea indirect detection by PPy-based amperometric biosensor. (b) Chronoamperometric test using Pt/PPy/poly(5-NH2 -1-NAP) electrode, without urease. Each urea addition corresponds to a 0.6 mmol L−1 increase in its concentration; 0.1 mol L−1 borate buffer (pH 10); E = 0.35 V.
2.3. Electrodeposition of monomers Pyrrole polymerization was performed potentiostatically at 0.7 V vs. Ag/AgCl/KClsat from an aqueous solution containing 50 mmol L−1 monomer and 25 mmol L−1 DBSA (dopant). The amount of PPy deposited over the electrode can be controlled by choosing proper deposition times, in order to obtain charge densities of PPy in the range 0.5–3.0 C cm−2 . Poly(5-NH2 -1-NAP) was electrochemically formed over predeposited PPy films via cyclic voltammetry (50 cycles, at 50 mV s−1 ) from an aqueous solution containing 5.0 mmol L−1 monomer and 1.0 mol L−1 HCl, according to the procedure described elsewhere [22]. The potential range of the two first cycles is 0–0.9 V, to lead to radical cations formation which will start the electropolymerization. From the third cycle onwards, the upper anodic potential was diminished to 0.7 V, to avoid degradation products production. 2.4. Ammonia amperometric detection
Scheme 1. Poly(5-amino-1-naphtol) structure.
The amperometric response of the sensor containing only PPy towards ammonia additions was verified in 0.1 mol L−1 borate buffer (pH 10.0) under continuous stirring. These tests were performed via chronoamperometry, at potentials ranging from 0.3 to
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0.4 V, and with charge densities of PPy between 0.5 and 3.0 C cm−2 . Amounts of ammonium ions solution with well-defined concentrations were periodically added to the system. All measurements were performed at 25 ± 1 ◦ C. 2.5. Urea biosensors 2.5.1. Immobilization of urease onto polymeric matrix The effect of different methods of urease immobilization onto the Pt electrode modified by 2.5 C cm−2 PPy and poly(5-NH2 -1NAP), 50 cycles was analyzed. The four methods studied were: 2.5.1.1. Adsorption. In these experiments, the electrode containing both polymers was placed in a 0.1 mg mL−1 urease aqueous solution for 3 h, at 4 ◦ C. 2.5.1.2. Entrapment. In this case, 10 L of 50 mg mL−1 urease solution prepared in 0.1 mol L−1 borate buffer pH 10 were deposited on the Pt/PPy/poly(5-NH2 -1-NAP) electrode, which was kept at 4 ◦ C until complete solvent evaporation. Thus, the amount of urease deposited onto this electrode is 0.5 mg. Then, 10 L of cellulose acetate 1% w/w in acetone was added onto the electrode. It was again kept at 4 ◦ C and after complete solvent evaporation the amperometric tests were performed.
3. Results and discussion 3.1. Optimization of ammonia sensor The first step for developing urea biosensors presented in this work is to choose the optimum working parameters for ammonia detection, once urea analysis in this case is directly dependent on ammonia sensing, as shown in Fig. 1a. Urea detection is indirect, via ammonia, and a first crucial result corroborating this mechanism is shown in Fig. 1b. In this experiment, a conventional amperometric test with urea additions is performed with an electrode containing both PPy and poly(5-NH2 -1-NAP), but no urease. No amperometric response was observed due to the fact that ammonia is not produced and the bipolymeric layer is sensitive to ammonia but not to urea. This result also proves that urea hydrolysis is a consequence of urease catalytic activity rather than a possible chemical decomposition because of the high pH employed. So, to optimize ammonia sensor parameters, the influence of both PPy charge density and working potential on the sensitivity towards ammonia was studied. Fig. 2a shows that sensitivity increases with PPy charge density from 0.5 up to 2.5 C cm−2 . This is an important result, which is in agreement with the fact that
2.5.1.3. Cross-linking. Here, 10 L of a freshly prepared solution containing 2.5 mg mL −1 urease, 5.0 mg mL−1 BSA and glutaraldehyde 0.16% v/v in 0.1 mol L−1 borate buffer pH 10 were casted onto the Pt/PPy/poly(5-NH2 -1-NAP) electrode. The mass of urease deposited using this procedure is then 0.05 mg. Then the modified electrode was kept at 4 ◦ C until the solvent was completely evaporated, and subsequently the amperometric tests were carried out. 2.5.1.4. Covalent attachment to poly(5-NH2 -1-NAP). The electrode containing both polymers was dipped in a 100 mg mL−1 cyanuric chloride ethanolic solution, under stirring, for 6 h. Then the electrode was rinsed with water and put into a 50 mg mL −1 urease solution prepared in 0.1 mol L−1 borate buffer (pH 10), also under stirring, for 15 h. To determine the mass of urease covalently attached to poly(5-NH2 -1-NAP), electrochemical quartz crystal microbalance (EQCM) experiments were carried out using a 6.0 MHz AT-cut overtone polished piezoelectric quartz crystal (Valpey–Fischer) with 25 mm diameter and 0.32 cm2 piezoactive electrode (gold) area (integral sensitive factor 6.45 × 107 cm2 Hz g−1 ). Frequency shifts were measured using a Stanford Research System model SR620 instrument connected to an oscillating circuit. The piezoactive gold electrode was modified with PPy and poly(5-NH2 -1-NAP) in the same charge densities used in all biosensors. The frequency shift between the electrode before (PPy/poly(5-NH2 1-NAP)/cyanuric chloride) and after enzyme linkage (PPy/poly(5NH2 -1-NAP)/cyanuric chloride/urease) enabled the quantification of urease chemically attached as being equal to 2.3 g. 2.5.2. Urea amperometric detection Urea amperometric detection was performed at 0.35 V vs. Ag/AgCl/KClsat , in 0.1 mol L−1 borate buffer solution for pHs ranging between 8 and 10, or in a 50 mmol L−1 phosphate buffer + 50 mmol L−1 KCl solution (pH 7.1). The working electrode consisted of (Pt/PPy 2.5 C cm−2 /poly(5-NH2 -1-NAP) 50 cycles) modified with urease in different forms, as described earlier. The electrolyte was continuously stirred and amounts of urea solution with well-defined concentration were periodically added to the system. All measurements were performed at 25 ± 1 ◦ C, and they were all repeated at least twice.
Fig. 2. Sensitivity of NH3 detection in 0.1 mol L−1 borate buffer (pH 10) as a function of: (a) charge density of poly(pyrrole), E = 0.35 V; (b) working potential, QPPy = 2.5 C cm−2 . Insert: cyclic voltamogram of 2.5 C cm−2 PPy in borate buffer 0.1 mol L−1 , pH 10.0, v = 50 mV s−1 . All experiments were performed in triplicate.
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PPy is indeed the transducer of the sensor, once the response is directly related to its amount present in the device. However, with 3.0 C cm−2 the response diminishes, just because this charge density leads to a thicker and non-adherent film, easily removed by a water jet. Charge densities lower than 0.5 C cm−2 were also studied, but negligible responses were obtained because those quantities were too low and insufficient to adequately act as the reaction transducer. Fig. 2b shows the dependence of sensitivity on the working potential, and it can be seen that the higher value was obtained at 0.35 V. This result can be confirmed by the cyclic voltamogram of 2.5 C cm−2 PPy film, shown as an insert in Fig. 2(b). In this voltamogram, the anodic potential peak is 0.35 V, showing the maximum electroactivity of PPy in this region. Thus, the parameters chosen for the development of urea biosensors were 2.5 C cm−2 of PPy and 0.35 V as working potential. The pH in which (bio)sensing is done must be at least 10 because it was observed that performing the tests at pHs below this value (7–9), amperometric response is much smaller, when not inexistent. This is in agreement with the fact that in those lower pHs, any or only a small amount of NH3 is formed (in these cases NH4 + is predominant, and it is not electroactive), providing very low current increases. So, all amperometric tests presented in this paper were performed at pH 10 (borate buffer medium, 0.1 mol L−1 ), although urease optimum working pH is found to be around 7.4. Results obtained at pH 10.0 show that urease activity is retained even at this high pH, enabling its application. 3.2. Interferents As will be shown later, poly(5-NH2 -1-NAP), which is also a conducting polymer, is essential for one of the enzyme immobilization methods (chemical attachment) employed. In that sense, to adequately compare the intricate contributions of every method, this polymer was used in the preparation of all of them. One of the benefits of using this polymer is that it drastically reduces the amperometric signal of interferents, such as ascorbic and uric acids, as shown in Fig. 3. This is an important feature to consider when dealing with clinical samples, in which those substances are present. In this case, the free –OH groups present in poly(5-NH2 -1NAP) chain provide a high electronic density to the polymer, acting
Fig. 3. Amperometric detection of NH3 , l-ascorbic acid and uric acid by sensors containing (grey) only PPy 2.5 C cm−2 ; (black) PPy 2.5 C cm−2 + poly(5-NH2 -1-NAP).
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as a electrostatic repulsion layer for the anionic interferents cited [20,22–23]. 3.3. Urea biosensors The development of all urea biosensors described in this work was based on the following assumptions: (a) ammonia sensor (which is the platform for urea biosensors) parameters must be chosen in order to achieve higher sensitivities, as discussed earlier; (b) the immobilization methods employed to attach urease to the polymeric matrix aim to achieve not only higher sensitivities but also longer stability. To do so and to compare the benefits of some kinds of immobilization methods, the incorporation of urease into the sensor was carried out through four different methods: (a) adsorption, (b) entrapment in a physical barrier of cellulose acetate, (c) cross-linking and (d) chemical attachment. The results are discussed in the next items, and it is important to emphasize that sensitivity values should be compared taking into account the amount of urease present in each biosensor; so that, the efficiency of each method for immobilizing the enzyme will also be discussed. 3.3.1. Immobilization methods 3.3.1.1. Adsorption. The first and simplest attempt to immobilize urease onto the electrode was through spontaneous adsorption, which occurs when species or regions of the species present some kind of attraction, for example Van der Waals or electrostatic attractions. The intention here was that some hydrophobic domains of urease could interact with both polymers, or even that urease could be somehow entrapped in the porous polymeric matrix. This kind of immobilization did not provide amperometric response towards urea, probably due to the fact that almost no urease could be immobilized through this method, because poly(5-NH2 -1-NAP) presents a high electronic density (free –OH groups) and urease is negatively charged at pH 10 (its isoelectric point is between 5.0 and 5.2 [25]), causing an electrostatic repulsion between them. 3.3.1.2. Entrapment in physical barrier of cellulose acetate. The second enzyme immobilization method consisted of the entrapment of this biomolecule inside a cellulose acetate polymeric network, aiming, for example, to reduce loss by leaching of the enzyme, one of the major problems found in this area of research. One advantage of this method is that conformational changes in the enzyme structure seldom happen, preserving its biological activity [10,24]. Fig. 4a presents the analytical curve related to this test, and the amperometric test is shown as an insert. The sensitivity towards urea in this case is low, 0.29 A cm−2 mmol−1 L, considering the amount of urease present in this biosensor (0.5 mg). This is probably due to the formation of a physical barrier of cellulose acetate that hinders urea diffusion, making the approach of this analyte to enzyme active sites more difficult. In this case, just a small percentage of urea is converted to ammonia, providing only a small current increase per total amount of urea added, causing low detection sensitivity. On the other hand, the linear fit adjusts very well to the experimental data, and the linear response range is wide (up to 40 mmol L−1 ), considering that urea concentration normally found in clinical samples (serum) is in the range 1.3–3.5 mmol L−1 [1]. 3.3.1.3. Cross-linking. Cross-linkages formation between glutaraldehyde and urease aims reducing enzyme, once there is a covalent bond formed between aldehyde end groups of glutaraldehyde and amino groups present in urease. Another advantage of using this immobilization method is that the interpenetrating network of urease-glutaraldehyde formed over the polymeric
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100 mmol L−1 ). A deviation is observed for urea concentrations above this value, and this behavior is attributed to poly(pyrrole) saturation [21] and also to enzymatic kinetics contribution. The biosensor used for the amperometric test showed in Fig. 4b was kept for 6 days in 0.1 mol L−1 borate buffer (pH 10.0), at 4 ◦ C. Then, urea amperometric detection was repeated with this electrode, and the result was a reduction in the sensitivity from 0.04 to 0.01 A cm−2 mmol−1 L. This reduction may be attributed to the fact that although there are linkages between urease and glutaraldehyde, there are no bonds between this network and the polymeric substrate. In that sense, this film loses adherence from the substrate, causing the leaching of a great amount of biomolecules and the consequent biosensor’s activity reduction.
Fig. 4. Analytical curves of the sensors containing both conducting polymers and urease incorporated via (a) entrapment; (b) cross-linking and (c) chemical attachment. The inserts show the amplification of the correspondent amperometric curves, in which each urea addition corresponds to a 0.6 mmol L−1 increase in its concentration, at 0.35 V, under usual conditions.
matrix provides easier urea diffusion until the enzyme active center, improving biosensor’s sensitivity. Fig. 4b shows the analytical curve obtained with this biosensor, and the correlated amperometric test is shown as an insert. This kind of immobilization method provided sensitivities lower (S = 0.04 A cm−2 mmol−1 L) than that obtained with the biosensor prepared via enzyme entrapment; but if one considers the mass of urease immobilized here (ten times smaller), the result is promising from the immobilization efficiency point of view. On the other hand, in Fig. 4b can be observed that the linear fit does not adjust to the experimental curve as well as in the last case. The linear concentration range is very wide (until
3.3.1.4. Covalent attachment. The preparation of this biosensor consisted of, in a first step, chemically attaching urease to poly(5NH2 -1-NAP) by means of an anchorage reaction. Cyanuric chloride was employed as anchor molecule, which attaches to poly(5-NH2 1-NAP) via its free –OH groups, so that urease can be covalently attached to the anchor by–NH2 groups present in the enzyme. The second step consisted of repeating the same procedure done with the biosensor in which immobilization was performed by cross-linking. This approach was chosen because using this biosensor’s design (containing both immobilization methods), if there is any improvement in this biosensor’s performance when compared with the last one (cross-linking), it would be strictly attributed to the chemical attachment of urease to the polymeric matrix, since it is the only difference between them. Here, the total amount of urease loaded onto the electrode is the sum of both contributions: chemical attachment of urease to poly(5-NH2 -1-NAP) (2.3 g) and cross-linking (0.05 mg). The biosensors described were tested in the same day of preparation and also after 6, 15 and 20 days. Between the measurements the electrodes were stored at 4 ◦ C in 0.1 mol L−1 borate buffer (pH 10). It was observed that sensitivity towards urea increased with time until the 15th day, reaching the maximum value of 0.50 A cm−2 mmol−1 L (Table 1). Fig. 4c shows the results concerning the experiment performed at the first day. In a first attempt to explain this behavior, the improvement of biosensor’s sensitivity was attributed to the continuous formation of bonds between cyanuric chloride and unbound urease, present in the surface of the electrode (due to glutaraldehyde-urease layer). However, experiments performed with a biosensor in which the only immobilization method used was the covalent attachment, without the addition of glutaraldehyde (the cross-linker), showed the same improvement in sensitivity with time. So, this improvement must be due to a better accommodation of the enzyme in the biosensor in order to expose its active site to the analyte. Besides, when comparing the sensitivities towards urea detection of these biosensors (chemical attachment of the enzyme) with those presented by others analyzed in this work, as described in Table 1 and in the second column of Table 2, it can be also observed that (1) the sensitivity of biosensors prepared by covalent enzyme immobilization is much higher than those prepared via entrapment and cross-linking, and the immobilization via entrapment led to sensitivities slightly below those obtained with cross-linking method; (2) from the 6th to the 15th day, the sensitivity of biosensors prepared via covalent attachment reach a stable value around 0.50 A cm−2 mmol−1 L, while all other biosensors have practically completely lost all their ability to detect urea. These biosensors present no more appreciable amperometric detection of urea only after the 20th day of preparation. Furthermore, from cross-linking to covalent attachment, there is a 1.05 increase in the mass of urease present in the biosensor, while the associated increase in sensitivity was 4.5 times. This result shows that the method of covalently attaching urease to the polymeric substrate provides a great gain
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Table 1 Mass of urease immobilized in each method and sensitivity of the different biosensors as function of time. Enzyme immobilization method
Mass of urease (mg)
First day S (A cm−2 mmol−1 L)
Sixth day S (A cm−2 mmol−1 L)
Fifteenth day S (A cm−2 mmol−1 L)
Adsorption Physical barrier Cross-linking Covalent attachment
– 5.00 × 10−1 5.00 × 10−2 5.23 × 10−2
– 0.29 0.04 0.18
– – 0.01 0.47
– – – 0.50
Table 2 Biosensors’ analytical parameters (sensitivity obtained at the first day normalized by the mass of urease, apparent Michaelis–Menten constant, saturation current and linear concentration range). Enzyme immobilization method
Normalized sensitivity (A cm−2 mmol−1 L mg−1 )
KMapp (mmol L−1 )
Imax (A cm−2 )
Linear concentration range (mol L−1 )
Adsorption Physical barrier Cross-linking Covalent attachment
– 0.58 0.80 3.44
– 131 4.6 137
– 43.0 0.3 19.0
– 0.22 × 10−3 –0.04 0.34 × 10−3 –0.10 0.58 × 10−3 –0.04
in sensitivity not only due to the slightly larger amount of enzyme, but to the intricate gain in conformational stability obtained when employing this new preparation method. The relative stability of biosensors presented in this item, in comparison with the others, is due to the effective chemical bonding between urease and poly(5-NH2 -1-NAP), not allowing the enzymes covalently attached to the polymeric substrate to be released into the solution. Besides, the close enzyme–substrate contact, through covalent bonding, may promote faster detection kinetics, once the electrochemically detectable species (ammonia) is produced closer to the transducer, reducing diffusion resistance, and thus increasing sensitivity and diminishing response times. Table 2 presents a summary of the analytical parameters obtained with the four biosensors studied, regarding the first day of testing. Sensitivity towards urea normalized by the total amount of immobilized urease, apparent Michaelis–Menten constant (KMapp ), saturated current (Imax ) and linear concentration range are depicted. KMapp and Imax were calculated from the linear part of the calibration curve, using the electrochemical version of Lineweaver–Burke equation (1/Iss = 1/Imax + KMapp /(Imax c), where Iss stands for stationary state current). It can be inferred from Table 2 that KMapp calculated for biosensors prepared using physical barrier and covalent attachment have approximately the same values, even if the amount of urease is about ten times smaller in the second case. This indicates an improvement in detection kinetics obtained with the covalent attachment of the enzyme to the transducer. From the comparison among some amperometric biosensors described in the literature and those containing urease chemically immobilized, presented in this work, it is possible to infer that the biosensors discussed here present a much extended linear concentration range and longer lifetimes than those obtained by the authors of references [13] (0.05–0.25 mmol L−1 ) and [14] (1.67–75.15 mol L−1 and 2 weeks of stability), in which urease is inserted in the PPy matrix concomitantly with polymerization of Py monomer. The biosensor analyzed in reference [26] is based on a urease membrane coupled with H+ -sensitive hydrazine electro-oxidation, and presented linear range (0.8–35.0 mmol L−1 ) a bit larger than the last one, probably due to the greater amount of enzyme present in the membrane covering the platinum electrode, as in the case of the adsorption of urease onto poly(vinylferrocenium) redox polymer [3] (1.0–250.0 mol L−1 ), which was performed electrostatically by anion-exchange, once the polymer is positively charged (PVF+ ) and the enzyme negatively (pH kept above the isoelectric point). In this last case, the 29 days stability is possible also due to the electrostatic interaction of species containing opposite charges (urease and the polymer). Rajesh et al. [1] developed a biosensor using chemical immobilization of urease via carbodimide coupling reaction between urease and the amino
groups of a pyrrole copolymer. This biosensor’s saturation current occurs at concentrations as low as 5.02 mmol L−1 . The sensitivity obtained in this case was 0.47 A cm−2 mmol L−1 , which is approximately the same value achieved with the biosensor prepared by covalent attachment of urease presented in this paper. 4. Conclusions This work outlined the results obtained with new biosensors designed for indirect detection of urea, via ammonia. The biosensors substrate consisted in a bipolymeric matrix of electrodeposited poly(pyrrole) and poly(5-amine-1-naphtol), the former being the electrochemical transducer of the amperometric reaction and the last playing the role of reducing interference signals of species like ascorbic and uric acids, improving selectivity. The immobilization of urease onto the substrate was performed through four different methods, and better results in terms of higher sensitivities towards urea detection and also longer biosensor’s lifetime were obtained with chemical attachment of the enzyme to poly(5-amine-1naphtol), due to an increased stability conferred by the chemical bond. On the other hand, all biosensors analyzed presented an extended linear concentration range (up to 100 mmol L−1 ) and also low detection limits (0.22–0.58 mmol L−1 ), considering their possible application in clinical samples, in which typical concentrations are found to be between 1.3 and 3.5 mmol L−1 [1]. The biosensors presented in this work have not been reported in literature before, neither this systematic comparison between the performances obtained when applying different enzyme immobilization methods to the same base sensor. Acknowledgements M.P.M. and S.I.C.T. gratefully acknowledge Brazilian Agencies FAPESP (06/53093-2 and 03/10015-3) and CNPq for financial support. References [1] V. Rajesh, W. Bisht, K. Takashima, Kaneto, An amperometric urea biosensor based on covalent immobilization of urease onto an electrochemically prepared copolymer poly (N-3-aminopropyl pyrrole-co-pyrrole) film, Biomaterials 26 (2005) 3683–3690. [2] R. Sahney, S. Anand, B.K. Puri, A.K. Srivastava, A comparative study of immobilization techniques for urease on glass-pH-electrode and its application in urea detection in blood serum, Anal. Chim. Acta 578 (2006) 156–161. [3] F. Kuralay, H. Özyörük, A. Yildiz, Amperometric enzyme electrode for urea determination using immobilized urease in poly(vinylferrocenium) film, Sens. Actuators, B, Chem. 114 (2006) 500–506. [4] S. Anne, J.R. Nicolo, M. Claude, C. Serge, A miniaturized urea sensor based on the integration of both ammonium based urea enzyme field effect transistor and a reference field effect transistor in a single chip, Talanta 50 (1999) 219–226.
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[5] G.G. Guilbault, M. Trap, A specific enzyme electrode for urea, Anal. Chim. Acta 73 (1974) 355–365. [6] G.G. Guilbault, M. Trap, Enzyme electrode based on the use of a carbon dioxide sensor, Anal. Chem. 44 (1972) 2161–2165. [7] P. Bataillard, E. Steffgen, S. Haemmerli, A. Manz, H.M. Widmer, An integrated silicon thermopile as biosensor for the thermal monitoring of glucose, urea and penicillin, Biosens. Bioelectron. 8 (1993) 89–98. [8] Y. Kawabata, H. Sugamoto, T. Imasaka, Micro-optrode for urea using an ammonium ion-sensitive membrane covered with a urease-immobilized membrane, Anal. Chim. Acta 283 (1993) 689–694. [9] H. Mana, U. Spohn, Selective flow injection procedures for the determination of nitrogen containing analytes by gasdialytic-fluorimetric detection of enzymatically generated ammonia, Anal. Chim. Acta 325 (1996) 93–104. [10] B. Lakard, G. Herlem, S. Lakard, A. Antoniou, B. Fahys, Urea potentiometric biosensor based on modified electrodes with urease immobilized on polyethylenimine films, Biosens. Bioelectron. 19 (2004) 1641–1647. [11] S.B. Adeloju, S.J. Shaw, G.G. Wallace, Polypyrrole-based potentiometric biosensor for urea part 2. analytical optimization, Anal. Chim. Acta 281 (1993) 621–627. [12] F. Kuralay, H. Özyörük, A. Yildiz, Potentiometric enzyme electrode for urea determination using immobilized urease in poly(vinylferrocenium) film, Sens. Actuators, B, Chem. 109 (2005) 194–199. [13] S.B. Adeloju, S.J. Shaw, G.G. Wallace, Polypyrrole-based amperometric flow injection biosensor for urea, Anal. Chim. Acta 323 (1996) 107–113. [14] S.B. Adeloju, S.J. Shaw, G.G. Wallace, Pulsed-amperometric detection of urea in blood samples on a conducting polypyrrole–urease biosensor, Anal. Chim. Acta 341 (1997) 155–160. [15] A. Ramanavicius, A. Ramanaviciene, A. Malinauskas, Electrochemical sensors based on conducting polymer—polypyrrole, Electrochim. Acta 51 (2006) 6025–6037. [16] S.M. Chung, W. Paik, I.H. Yeo, A study on the initial growth of polypyrrole on a gold electrode by electrochemical quartz crystal microbalance, Synt. Metals 84 (1997) 155–156. [17] G. Appel, O. Böhme, R. Mikalo, D. Schmeier, The polaron and bipolaron contributions to the electronic structure of polypyrrole films, Chem. Phys. Lett. 313 (1999) 411–415. [18] A. Kaynak, Effect of synthesis parameters on the surface morphology of conducting polypyrrole films, Mat. Res. Bull. 32 (1997) 271–285. [19] R.G. Davidson, L.C. Hammond, T.G. Turner, A.R. Wilson, An electron and X-ray diffraction study of conducting polypyrrole/dodecyl sulfate, Synt. Metals 81 (1996) 1–4. [20] M. Vidotti, L.H. Dall’Antonia, E.P. Cintra, S.I. Córdoba de Torresi, Reduction of interference signal of ascorbate and urate in poly(pyrrole)-based ammonia sensors in aqueous solutions, Electrochim. Acta 49 (2004) 3665–3670. [21] L.H. Dall’Antonia, M.E. Vidotti, S.I. Córdoba de Torresi, R.M. Torresi, A new sensor for ammonia determination based on polypyrrole films doped with dodecylbenzenesulfonate (DBSA) ions, Electroanalysis 14 (2002) 1577–1586.
[22] E.P. Cintra, S.I. Córdoba de Torresi, Resonant Raman spectroscopy as a tool for determining the formation of a ladder structure in electropolymerized poly(5amino-1-naphthol), J. Electroanal. Chem. 518 (2002) 33–40. [23] E.P. Cintra, R.M. Torresi, G. Louarn, S.I. Córdoba de Torresi, Electronic and ionic exchange in poly(5-amino-1-naphthol) in acid aqueous solution, Electrochim. Acta 49 (2004) 1409–1415. [24] M. Singh, N. Verma, A.K. Garg, N. Redhu, Urea biosensors, Sens. Actuators, B, Chem. 134 (2008) 345–351. [25] J.B. Sumner, D.B. Hand, The isoelectric point of crystalline urease, J. Am. Chem. Soc. 51 (1929) 1255–1260. [26] D. Kirstein, L. Kirstein, F. Scheller, Enzyme electrode for urea with amperometric indication: part I—basic principle, Biosensors 1 (1985) 117–130.
Biographies Mariana P. Massafera was born in São Bernardo do Campo, Brazil, on 7 May 1984. She received her bachelor degree in chemistry from Chemistry Institute at São Carlos–Universidade de São Paulo in 2005. She has been a PhD student in Chemistry Institute–Universidade de São Paulo since then. Her research interests include biosensors, conducting polymers, nanotechnology and carbon nanotubes composites.
Susana I. Córdoba de Torresi graduated in physical chemistry at Universidad Nacional de Córdoba, Argentina (1984), where she also concluded PhD in chemistry with emphasis in Electrochemistry in 1988. She was post-doc of Université de Paris VI, France (1989–1990), Physics Institute of UNICAMP, Brazil (1991–1993) and Chemistry Department of UFSCar, Brazil (1994–1995). Presently, she is full professor of the Chemistry Institute, Universidade de São Paulo since 1996. Her interests are in the fields of electrochemical devices, nanotechnology and tailoring of new materials.
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Sensors and Actuators B: Chemical journal homepage: www.elsevier.com/locate/snb
Urea amperometric biosensors based on a multifunctional bipolymeric layer: Comparing enzyme immobilization methods Mariana P. Massafera, Susana I. Córdoba de Torresi ∗ Instituto de Química, Universidade de São Paulo, C.P. 26077, 05513-970 São Paulo (SP), Brazil
a r t i c l e
i n f o
Article history: Received 29 July 2008 Received in revised form 29 January 2009 Accepted 5 February 2009 Available online 20 February 2009 Keywords: Biosensors Urease Poly(pyrrole) Poly(5-amino-1-naphthol) Urea Ammonia
a b s t r a c t This paper outlines the results obtained with biosensors designed for urea amperometric detection. The incorporation of urease into a bipolymeric substrate consisting of poly(pyrrole) and poly(5-amino-1naphthol) was performed through four different approaches: direct adsorption, entrapment in cellulose acetate layer, cross-linking with glutaraldehyde, and also covalent attachment to the polymeric matrix. Poly(pyrrole) acts as amperometric transducer in these biosensors, while poly(5-amino-1-naphthol) drastically reduces the interference signal of agents such as ascorbic and uric acids. The biosensors containing urease covalently attached to the substrate provided interesting results in terms of sensitivity towards urea (0.50 A cm−2 mmol−1 L), lifetime (20 days) and short response times, due to the enzyme immobilization method used. All biosensors analyzed showed also a wide linear concentration range (up to 100 mmol L−1 ) and low detection limits (0.22–0.58 mmol L−1 ). © 2009 Elsevier B.V. All rights reserved.
1. Introduction In human organism, as urea is a final product of protein metabolism, its accumulation above critical levels must be carefully determined. So, urea monitoring is of great interest in clinical and biomedical analysis because this biomolecule enables the diagnosis of some diseases such as kidney and liver malfunction, and also make possible to predict the nature and causes of diabetes [1–3]. Urea is also related to other processes such as soil fertilization, in which it acts as a nitrogen source for plants. For the purpose of urea sensing, various procedures are applied and interesting results have been obtained with enzyme-based sensors. The incorporation of a biological moiety (enzymes, cells, antigens, DNA) to a conventional physico-chemical transducer characterizes a biosensor, and the presence of the biomolecule provides a high specificity to biosensors. The enzyme used in urea biosensors, urease, catalyses the hydrolysis of urea producing ammonium and bicarbonate ions, causing a pH increase in the aqueous reaction medium, as shown in the following reaction: urease
NH2 CONH2 + 3H2 O −→ 2NH4 + + HCO3 − + OH−
(1)
Considering this reaction, the determination of urea can be indirectly done by using specific transducers that detect ammo-
∗ Corresponding author. Tel.: +55 11 30912350; fax: +55 11 37855579. E-mail address: [email protected] (S.I.C.d. Torresi). 0925-4005/$ – see front matter © 2009 Elsevier B.V. All rights reserved. doi:10.1016/j.snb.2009.02.013
nium ions [4], ammonia gas [5], carbon dioxide [6], or pH changes [2–3]. Alternatively, techniques as calorimetric [7] and fluorimetric analysis [8–9] are also available, but one disadvantage of these methods is that in most cases, the sample must be diluted and pretreated before analysis. However, biosensors containing electrochemical transducers are the most used and in these cases urea is detected via either potentiometric [10–12] or amperometric methods [1,3,13–14]. This last one has attracted a lot of attention due to both its effectiveness and simplicity of preparation. Regarding electrochemical biosensors, conducting polymers can be employed as very suitable transducers. An outstanding characteristic of these polymers is their ability to reversibly alternate between the conducting and insulating electronic states. This alternation occurs because conducting polymers present in their structure sequences of conjugated -bonds, changing from insulating to conductors by an oxidation or reduction process, in which incorporation of ions can take place (doping). Many other properties of these special polymers support their application in biosensors, but mainly their ease of preparation (direct deposition of a polymeric film onto an electrode can be performed by electrochemical oxidation of monomers) and other controlled parameters (film thickness, kind of dopant). Among conducting polymers, poly(pyrrole) (PPy) plays an important role because it presents extremely favorable characteristics, as for example its easy synthesis, good electrochemical reversibility, excellent stability, besides presenting a good electroactivity in a large pH range [15–19]. Another important feature of conducting polymers is that some of them can act as selective layers, reducing the amperometric
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signal of interferents as ascorbic and uric acids, species which are often present in clinical samples. This is the case of poly(5-amino-1naphthol) (poly(5-NH2 -1-NAP), Scheme 1), a polymer that was used in addition to PPy in the development of ammonia amperometric sensors [20–21]. Poly(5-NH2 -1-NAP) presents a “polyaniline-like” structure with free –OH groups [22–23] that can undergo a chemical reaction to provide a covalent linkage between the polymeric chain and an enzyme, for example urease, as it will be shown in this work. Covalent attachment of the biomolecule to the substrate is one of the most elegant immobilization methods available, but others as adsorption, entrapment and cross-linking are often used [24]. As will be discussed later, the covalent attachment of the biomolecule to the transducer presents the advantage of enhancing biosensors stability, once one of the major problems of biosensors preparation is the leaching of the biomolecule out of the substrate. The purpose of this work is to provide a systematic comparison among urea amperometric biosensors prepared through those four immobilization methods explained above, using as biological recognition element the enzyme urease. The biosensors consist of a bipolymeric substrate of PPy and poly(5-NH2 -1-NAP), the first acting as electrochemical transducer for ammonia detection (as shown in Fig. 1a), substance that is produced during enzymatic hydrolysis of urea in alkaline medium (reaction (1)). The second plays a double role once it is necessary to covalently attach urease to the substrate and also to reduce interferents signal. 2. Experimental 2.1. Reagents Pyrrole (Aldrich, 99%) was purified by distillation and kept in refrigerator under N2 atmosphere, in the dark. Dodecilbenzenesulfonate, sodium salt (DBSA), 5-amino-1-naphtol (5-NH2 -1-NAP) 97%, cellulose acetate, cyanuric chloride 99% (Aldrich), NaOH, HCl, H3 BO3 (Synth, analytical grade), glutaraldehyde 25% v/v (Merck), NH4 Cl (Mallinckrodt, analytical grade), l-ascorbic and uric acids, bovine serum albumin (BSA) 200 mg mL−1 and Urease (EC 3.5.1.5, type III from Jack Beans, 29,500 units g−1 ) (Sigma) were used as received. All the solutions were prepared using water purified by an Elga UHQ system. 2.2. Instrumental Monomers electropolymerization and chronoamperometric tests (both ammonia and urea detection) were performed in a three electrode conventional cell consisting of a working electrode (Pt/PPy for ammonia detection, and Pt/PPy/poly(5-NH2 -1-NAP) modified with urease in case of urea determination), a Pt foil as counter-electrode and Ag/AgCl/KClsat as the reference electrode. The electrochemical data were obtained via an Autolab equipment model PGSTAT 30 (Eco Chemie), interfaced with a computer for data acquisition.
Fig. 1. (a) Urea indirect detection by PPy-based amperometric biosensor. (b) Chronoamperometric test using Pt/PPy/poly(5-NH2 -1-NAP) electrode, without urease. Each urea addition corresponds to a 0.6 mmol L−1 increase in its concentration; 0.1 mol L−1 borate buffer (pH 10); E = 0.35 V.
2.3. Electrodeposition of monomers Pyrrole polymerization was performed potentiostatically at 0.7 V vs. Ag/AgCl/KClsat from an aqueous solution containing 50 mmol L−1 monomer and 25 mmol L−1 DBSA (dopant). The amount of PPy deposited over the electrode can be controlled by choosing proper deposition times, in order to obtain charge densities of PPy in the range 0.5–3.0 C cm−2 . Poly(5-NH2 -1-NAP) was electrochemically formed over predeposited PPy films via cyclic voltammetry (50 cycles, at 50 mV s−1 ) from an aqueous solution containing 5.0 mmol L−1 monomer and 1.0 mol L−1 HCl, according to the procedure described elsewhere [22]. The potential range of the two first cycles is 0–0.9 V, to lead to radical cations formation which will start the electropolymerization. From the third cycle onwards, the upper anodic potential was diminished to 0.7 V, to avoid degradation products production. 2.4. Ammonia amperometric detection
Scheme 1. Poly(5-amino-1-naphtol) structure.
The amperometric response of the sensor containing only PPy towards ammonia additions was verified in 0.1 mol L−1 borate buffer (pH 10.0) under continuous stirring. These tests were performed via chronoamperometry, at potentials ranging from 0.3 to
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0.4 V, and with charge densities of PPy between 0.5 and 3.0 C cm−2 . Amounts of ammonium ions solution with well-defined concentrations were periodically added to the system. All measurements were performed at 25 ± 1 ◦ C. 2.5. Urea biosensors 2.5.1. Immobilization of urease onto polymeric matrix The effect of different methods of urease immobilization onto the Pt electrode modified by 2.5 C cm−2 PPy and poly(5-NH2 -1NAP), 50 cycles was analyzed. The four methods studied were: 2.5.1.1. Adsorption. In these experiments, the electrode containing both polymers was placed in a 0.1 mg mL−1 urease aqueous solution for 3 h, at 4 ◦ C. 2.5.1.2. Entrapment. In this case, 10 L of 50 mg mL−1 urease solution prepared in 0.1 mol L−1 borate buffer pH 10 were deposited on the Pt/PPy/poly(5-NH2 -1-NAP) electrode, which was kept at 4 ◦ C until complete solvent evaporation. Thus, the amount of urease deposited onto this electrode is 0.5 mg. Then, 10 L of cellulose acetate 1% w/w in acetone was added onto the electrode. It was again kept at 4 ◦ C and after complete solvent evaporation the amperometric tests were performed.
3. Results and discussion 3.1. Optimization of ammonia sensor The first step for developing urea biosensors presented in this work is to choose the optimum working parameters for ammonia detection, once urea analysis in this case is directly dependent on ammonia sensing, as shown in Fig. 1a. Urea detection is indirect, via ammonia, and a first crucial result corroborating this mechanism is shown in Fig. 1b. In this experiment, a conventional amperometric test with urea additions is performed with an electrode containing both PPy and poly(5-NH2 -1-NAP), but no urease. No amperometric response was observed due to the fact that ammonia is not produced and the bipolymeric layer is sensitive to ammonia but not to urea. This result also proves that urea hydrolysis is a consequence of urease catalytic activity rather than a possible chemical decomposition because of the high pH employed. So, to optimize ammonia sensor parameters, the influence of both PPy charge density and working potential on the sensitivity towards ammonia was studied. Fig. 2a shows that sensitivity increases with PPy charge density from 0.5 up to 2.5 C cm−2 . This is an important result, which is in agreement with the fact that
2.5.1.3. Cross-linking. Here, 10 L of a freshly prepared solution containing 2.5 mg mL −1 urease, 5.0 mg mL−1 BSA and glutaraldehyde 0.16% v/v in 0.1 mol L−1 borate buffer pH 10 were casted onto the Pt/PPy/poly(5-NH2 -1-NAP) electrode. The mass of urease deposited using this procedure is then 0.05 mg. Then the modified electrode was kept at 4 ◦ C until the solvent was completely evaporated, and subsequently the amperometric tests were carried out. 2.5.1.4. Covalent attachment to poly(5-NH2 -1-NAP). The electrode containing both polymers was dipped in a 100 mg mL−1 cyanuric chloride ethanolic solution, under stirring, for 6 h. Then the electrode was rinsed with water and put into a 50 mg mL −1 urease solution prepared in 0.1 mol L−1 borate buffer (pH 10), also under stirring, for 15 h. To determine the mass of urease covalently attached to poly(5-NH2 -1-NAP), electrochemical quartz crystal microbalance (EQCM) experiments were carried out using a 6.0 MHz AT-cut overtone polished piezoelectric quartz crystal (Valpey–Fischer) with 25 mm diameter and 0.32 cm2 piezoactive electrode (gold) area (integral sensitive factor 6.45 × 107 cm2 Hz g−1 ). Frequency shifts were measured using a Stanford Research System model SR620 instrument connected to an oscillating circuit. The piezoactive gold electrode was modified with PPy and poly(5-NH2 -1-NAP) in the same charge densities used in all biosensors. The frequency shift between the electrode before (PPy/poly(5-NH2 1-NAP)/cyanuric chloride) and after enzyme linkage (PPy/poly(5NH2 -1-NAP)/cyanuric chloride/urease) enabled the quantification of urease chemically attached as being equal to 2.3 g. 2.5.2. Urea amperometric detection Urea amperometric detection was performed at 0.35 V vs. Ag/AgCl/KClsat , in 0.1 mol L−1 borate buffer solution for pHs ranging between 8 and 10, or in a 50 mmol L−1 phosphate buffer + 50 mmol L−1 KCl solution (pH 7.1). The working electrode consisted of (Pt/PPy 2.5 C cm−2 /poly(5-NH2 -1-NAP) 50 cycles) modified with urease in different forms, as described earlier. The electrolyte was continuously stirred and amounts of urea solution with well-defined concentration were periodically added to the system. All measurements were performed at 25 ± 1 ◦ C, and they were all repeated at least twice.
Fig. 2. Sensitivity of NH3 detection in 0.1 mol L−1 borate buffer (pH 10) as a function of: (a) charge density of poly(pyrrole), E = 0.35 V; (b) working potential, QPPy = 2.5 C cm−2 . Insert: cyclic voltamogram of 2.5 C cm−2 PPy in borate buffer 0.1 mol L−1 , pH 10.0, v = 50 mV s−1 . All experiments were performed in triplicate.
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PPy is indeed the transducer of the sensor, once the response is directly related to its amount present in the device. However, with 3.0 C cm−2 the response diminishes, just because this charge density leads to a thicker and non-adherent film, easily removed by a water jet. Charge densities lower than 0.5 C cm−2 were also studied, but negligible responses were obtained because those quantities were too low and insufficient to adequately act as the reaction transducer. Fig. 2b shows the dependence of sensitivity on the working potential, and it can be seen that the higher value was obtained at 0.35 V. This result can be confirmed by the cyclic voltamogram of 2.5 C cm−2 PPy film, shown as an insert in Fig. 2(b). In this voltamogram, the anodic potential peak is 0.35 V, showing the maximum electroactivity of PPy in this region. Thus, the parameters chosen for the development of urea biosensors were 2.5 C cm−2 of PPy and 0.35 V as working potential. The pH in which (bio)sensing is done must be at least 10 because it was observed that performing the tests at pHs below this value (7–9), amperometric response is much smaller, when not inexistent. This is in agreement with the fact that in those lower pHs, any or only a small amount of NH3 is formed (in these cases NH4 + is predominant, and it is not electroactive), providing very low current increases. So, all amperometric tests presented in this paper were performed at pH 10 (borate buffer medium, 0.1 mol L−1 ), although urease optimum working pH is found to be around 7.4. Results obtained at pH 10.0 show that urease activity is retained even at this high pH, enabling its application. 3.2. Interferents As will be shown later, poly(5-NH2 -1-NAP), which is also a conducting polymer, is essential for one of the enzyme immobilization methods (chemical attachment) employed. In that sense, to adequately compare the intricate contributions of every method, this polymer was used in the preparation of all of them. One of the benefits of using this polymer is that it drastically reduces the amperometric signal of interferents, such as ascorbic and uric acids, as shown in Fig. 3. This is an important feature to consider when dealing with clinical samples, in which those substances are present. In this case, the free –OH groups present in poly(5-NH2 -1NAP) chain provide a high electronic density to the polymer, acting
Fig. 3. Amperometric detection of NH3 , l-ascorbic acid and uric acid by sensors containing (grey) only PPy 2.5 C cm−2 ; (black) PPy 2.5 C cm−2 + poly(5-NH2 -1-NAP).
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as a electrostatic repulsion layer for the anionic interferents cited [20,22–23]. 3.3. Urea biosensors The development of all urea biosensors described in this work was based on the following assumptions: (a) ammonia sensor (which is the platform for urea biosensors) parameters must be chosen in order to achieve higher sensitivities, as discussed earlier; (b) the immobilization methods employed to attach urease to the polymeric matrix aim to achieve not only higher sensitivities but also longer stability. To do so and to compare the benefits of some kinds of immobilization methods, the incorporation of urease into the sensor was carried out through four different methods: (a) adsorption, (b) entrapment in a physical barrier of cellulose acetate, (c) cross-linking and (d) chemical attachment. The results are discussed in the next items, and it is important to emphasize that sensitivity values should be compared taking into account the amount of urease present in each biosensor; so that, the efficiency of each method for immobilizing the enzyme will also be discussed. 3.3.1. Immobilization methods 3.3.1.1. Adsorption. The first and simplest attempt to immobilize urease onto the electrode was through spontaneous adsorption, which occurs when species or regions of the species present some kind of attraction, for example Van der Waals or electrostatic attractions. The intention here was that some hydrophobic domains of urease could interact with both polymers, or even that urease could be somehow entrapped in the porous polymeric matrix. This kind of immobilization did not provide amperometric response towards urea, probably due to the fact that almost no urease could be immobilized through this method, because poly(5-NH2 -1-NAP) presents a high electronic density (free –OH groups) and urease is negatively charged at pH 10 (its isoelectric point is between 5.0 and 5.2 [25]), causing an electrostatic repulsion between them. 3.3.1.2. Entrapment in physical barrier of cellulose acetate. The second enzyme immobilization method consisted of the entrapment of this biomolecule inside a cellulose acetate polymeric network, aiming, for example, to reduce loss by leaching of the enzyme, one of the major problems found in this area of research. One advantage of this method is that conformational changes in the enzyme structure seldom happen, preserving its biological activity [10,24]. Fig. 4a presents the analytical curve related to this test, and the amperometric test is shown as an insert. The sensitivity towards urea in this case is low, 0.29 A cm−2 mmol−1 L, considering the amount of urease present in this biosensor (0.5 mg). This is probably due to the formation of a physical barrier of cellulose acetate that hinders urea diffusion, making the approach of this analyte to enzyme active sites more difficult. In this case, just a small percentage of urea is converted to ammonia, providing only a small current increase per total amount of urea added, causing low detection sensitivity. On the other hand, the linear fit adjusts very well to the experimental data, and the linear response range is wide (up to 40 mmol L−1 ), considering that urea concentration normally found in clinical samples (serum) is in the range 1.3–3.5 mmol L−1 [1]. 3.3.1.3. Cross-linking. Cross-linkages formation between glutaraldehyde and urease aims reducing enzyme, once there is a covalent bond formed between aldehyde end groups of glutaraldehyde and amino groups present in urease. Another advantage of using this immobilization method is that the interpenetrating network of urease-glutaraldehyde formed over the polymeric
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100 mmol L−1 ). A deviation is observed for urea concentrations above this value, and this behavior is attributed to poly(pyrrole) saturation [21] and also to enzymatic kinetics contribution. The biosensor used for the amperometric test showed in Fig. 4b was kept for 6 days in 0.1 mol L−1 borate buffer (pH 10.0), at 4 ◦ C. Then, urea amperometric detection was repeated with this electrode, and the result was a reduction in the sensitivity from 0.04 to 0.01 A cm−2 mmol−1 L. This reduction may be attributed to the fact that although there are linkages between urease and glutaraldehyde, there are no bonds between this network and the polymeric substrate. In that sense, this film loses adherence from the substrate, causing the leaching of a great amount of biomolecules and the consequent biosensor’s activity reduction.
Fig. 4. Analytical curves of the sensors containing both conducting polymers and urease incorporated via (a) entrapment; (b) cross-linking and (c) chemical attachment. The inserts show the amplification of the correspondent amperometric curves, in which each urea addition corresponds to a 0.6 mmol L−1 increase in its concentration, at 0.35 V, under usual conditions.
matrix provides easier urea diffusion until the enzyme active center, improving biosensor’s sensitivity. Fig. 4b shows the analytical curve obtained with this biosensor, and the correlated amperometric test is shown as an insert. This kind of immobilization method provided sensitivities lower (S = 0.04 A cm−2 mmol−1 L) than that obtained with the biosensor prepared via enzyme entrapment; but if one considers the mass of urease immobilized here (ten times smaller), the result is promising from the immobilization efficiency point of view. On the other hand, in Fig. 4b can be observed that the linear fit does not adjust to the experimental curve as well as in the last case. The linear concentration range is very wide (until
3.3.1.4. Covalent attachment. The preparation of this biosensor consisted of, in a first step, chemically attaching urease to poly(5NH2 -1-NAP) by means of an anchorage reaction. Cyanuric chloride was employed as anchor molecule, which attaches to poly(5-NH2 1-NAP) via its free –OH groups, so that urease can be covalently attached to the anchor by–NH2 groups present in the enzyme. The second step consisted of repeating the same procedure done with the biosensor in which immobilization was performed by cross-linking. This approach was chosen because using this biosensor’s design (containing both immobilization methods), if there is any improvement in this biosensor’s performance when compared with the last one (cross-linking), it would be strictly attributed to the chemical attachment of urease to the polymeric matrix, since it is the only difference between them. Here, the total amount of urease loaded onto the electrode is the sum of both contributions: chemical attachment of urease to poly(5-NH2 -1-NAP) (2.3 g) and cross-linking (0.05 mg). The biosensors described were tested in the same day of preparation and also after 6, 15 and 20 days. Between the measurements the electrodes were stored at 4 ◦ C in 0.1 mol L−1 borate buffer (pH 10). It was observed that sensitivity towards urea increased with time until the 15th day, reaching the maximum value of 0.50 A cm−2 mmol−1 L (Table 1). Fig. 4c shows the results concerning the experiment performed at the first day. In a first attempt to explain this behavior, the improvement of biosensor’s sensitivity was attributed to the continuous formation of bonds between cyanuric chloride and unbound urease, present in the surface of the electrode (due to glutaraldehyde-urease layer). However, experiments performed with a biosensor in which the only immobilization method used was the covalent attachment, without the addition of glutaraldehyde (the cross-linker), showed the same improvement in sensitivity with time. So, this improvement must be due to a better accommodation of the enzyme in the biosensor in order to expose its active site to the analyte. Besides, when comparing the sensitivities towards urea detection of these biosensors (chemical attachment of the enzyme) with those presented by others analyzed in this work, as described in Table 1 and in the second column of Table 2, it can be also observed that (1) the sensitivity of biosensors prepared by covalent enzyme immobilization is much higher than those prepared via entrapment and cross-linking, and the immobilization via entrapment led to sensitivities slightly below those obtained with cross-linking method; (2) from the 6th to the 15th day, the sensitivity of biosensors prepared via covalent attachment reach a stable value around 0.50 A cm−2 mmol−1 L, while all other biosensors have practically completely lost all their ability to detect urea. These biosensors present no more appreciable amperometric detection of urea only after the 20th day of preparation. Furthermore, from cross-linking to covalent attachment, there is a 1.05 increase in the mass of urease present in the biosensor, while the associated increase in sensitivity was 4.5 times. This result shows that the method of covalently attaching urease to the polymeric substrate provides a great gain
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Table 1 Mass of urease immobilized in each method and sensitivity of the different biosensors as function of time. Enzyme immobilization method
Mass of urease (mg)
First day S (A cm−2 mmol−1 L)
Sixth day S (A cm−2 mmol−1 L)
Fifteenth day S (A cm−2 mmol−1 L)
Adsorption Physical barrier Cross-linking Covalent attachment
– 5.00 × 10−1 5.00 × 10−2 5.23 × 10−2
– 0.29 0.04 0.18
– – 0.01 0.47
– – – 0.50
Table 2 Biosensors’ analytical parameters (sensitivity obtained at the first day normalized by the mass of urease, apparent Michaelis–Menten constant, saturation current and linear concentration range). Enzyme immobilization method
Normalized sensitivity (A cm−2 mmol−1 L mg−1 )
KMapp (mmol L−1 )
Imax (A cm−2 )
Linear concentration range (mol L−1 )
Adsorption Physical barrier Cross-linking Covalent attachment
– 0.58 0.80 3.44
– 131 4.6 137
– 43.0 0.3 19.0
– 0.22 × 10−3 –0.04 0.34 × 10−3 –0.10 0.58 × 10−3 –0.04
in sensitivity not only due to the slightly larger amount of enzyme, but to the intricate gain in conformational stability obtained when employing this new preparation method. The relative stability of biosensors presented in this item, in comparison with the others, is due to the effective chemical bonding between urease and poly(5-NH2 -1-NAP), not allowing the enzymes covalently attached to the polymeric substrate to be released into the solution. Besides, the close enzyme–substrate contact, through covalent bonding, may promote faster detection kinetics, once the electrochemically detectable species (ammonia) is produced closer to the transducer, reducing diffusion resistance, and thus increasing sensitivity and diminishing response times. Table 2 presents a summary of the analytical parameters obtained with the four biosensors studied, regarding the first day of testing. Sensitivity towards urea normalized by the total amount of immobilized urease, apparent Michaelis–Menten constant (KMapp ), saturated current (Imax ) and linear concentration range are depicted. KMapp and Imax were calculated from the linear part of the calibration curve, using the electrochemical version of Lineweaver–Burke equation (1/Iss = 1/Imax + KMapp /(Imax c), where Iss stands for stationary state current). It can be inferred from Table 2 that KMapp calculated for biosensors prepared using physical barrier and covalent attachment have approximately the same values, even if the amount of urease is about ten times smaller in the second case. This indicates an improvement in detection kinetics obtained with the covalent attachment of the enzyme to the transducer. From the comparison among some amperometric biosensors described in the literature and those containing urease chemically immobilized, presented in this work, it is possible to infer that the biosensors discussed here present a much extended linear concentration range and longer lifetimes than those obtained by the authors of references [13] (0.05–0.25 mmol L−1 ) and [14] (1.67–75.15 mol L−1 and 2 weeks of stability), in which urease is inserted in the PPy matrix concomitantly with polymerization of Py monomer. The biosensor analyzed in reference [26] is based on a urease membrane coupled with H+ -sensitive hydrazine electro-oxidation, and presented linear range (0.8–35.0 mmol L−1 ) a bit larger than the last one, probably due to the greater amount of enzyme present in the membrane covering the platinum electrode, as in the case of the adsorption of urease onto poly(vinylferrocenium) redox polymer [3] (1.0–250.0 mol L−1 ), which was performed electrostatically by anion-exchange, once the polymer is positively charged (PVF+ ) and the enzyme negatively (pH kept above the isoelectric point). In this last case, the 29 days stability is possible also due to the electrostatic interaction of species containing opposite charges (urease and the polymer). Rajesh et al. [1] developed a biosensor using chemical immobilization of urease via carbodimide coupling reaction between urease and the amino
groups of a pyrrole copolymer. This biosensor’s saturation current occurs at concentrations as low as 5.02 mmol L−1 . The sensitivity obtained in this case was 0.47 A cm−2 mmol L−1 , which is approximately the same value achieved with the biosensor prepared by covalent attachment of urease presented in this paper. 4. Conclusions This work outlined the results obtained with new biosensors designed for indirect detection of urea, via ammonia. The biosensors substrate consisted in a bipolymeric matrix of electrodeposited poly(pyrrole) and poly(5-amine-1-naphtol), the former being the electrochemical transducer of the amperometric reaction and the last playing the role of reducing interference signals of species like ascorbic and uric acids, improving selectivity. The immobilization of urease onto the substrate was performed through four different methods, and better results in terms of higher sensitivities towards urea detection and also longer biosensor’s lifetime were obtained with chemical attachment of the enzyme to poly(5-amine-1naphtol), due to an increased stability conferred by the chemical bond. On the other hand, all biosensors analyzed presented an extended linear concentration range (up to 100 mmol L−1 ) and also low detection limits (0.22–0.58 mmol L−1 ), considering their possible application in clinical samples, in which typical concentrations are found to be between 1.3 and 3.5 mmol L−1 [1]. The biosensors presented in this work have not been reported in literature before, neither this systematic comparison between the performances obtained when applying different enzyme immobilization methods to the same base sensor. Acknowledgements M.P.M. and S.I.C.T. gratefully acknowledge Brazilian Agencies FAPESP (06/53093-2 and 03/10015-3) and CNPq for financial support. References [1] V. Rajesh, W. Bisht, K. Takashima, Kaneto, An amperometric urea biosensor based on covalent immobilization of urease onto an electrochemically prepared copolymer poly (N-3-aminopropyl pyrrole-co-pyrrole) film, Biomaterials 26 (2005) 3683–3690. [2] R. Sahney, S. Anand, B.K. Puri, A.K. Srivastava, A comparative study of immobilization techniques for urease on glass-pH-electrode and its application in urea detection in blood serum, Anal. Chim. Acta 578 (2006) 156–161. [3] F. Kuralay, H. Özyörük, A. Yildiz, Amperometric enzyme electrode for urea determination using immobilized urease in poly(vinylferrocenium) film, Sens. Actuators, B, Chem. 114 (2006) 500–506. [4] S. Anne, J.R. Nicolo, M. Claude, C. Serge, A miniaturized urea sensor based on the integration of both ammonium based urea enzyme field effect transistor and a reference field effect transistor in a single chip, Talanta 50 (1999) 219–226.
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Biographies Mariana P. Massafera was born in São Bernardo do Campo, Brazil, on 7 May 1984. She received her bachelor degree in chemistry from Chemistry Institute at São Carlos–Universidade de São Paulo in 2005. She has been a PhD student in Chemistry Institute–Universidade de São Paulo since then. Her research interests include biosensors, conducting polymers, nanotechnology and carbon nanotubes composites.
Susana I. Córdoba de Torresi graduated in physical chemistry at Universidad Nacional de Córdoba, Argentina (1984), where she also concluded PhD in chemistry with emphasis in Electrochemistry in 1988. She was post-doc of Université de Paris VI, France (1989–1990), Physics Institute of UNICAMP, Brazil (1991–1993) and Chemistry Department of UFSCar, Brazil (1994–1995). Presently, she is full professor of the Chemistry Institute, Universidade de São Paulo since 1996. Her interests are in the fields of electrochemical devices, nanotechnology and tailoring of new materials.