Effects of self-assembled monolayers on amperometric glucose biosensors based on an organic–inorganic hybrid system

Sensors and Actuators B 168 (2012) 249–255

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Sensors and Actuators B: Chemical journal homepage: www.elsevier.com/locate/snb

Effects of self-assembled monolayers on amperometric glucose biosensors based on an organic–inorganic hybrid system Huihui Wang a , Hitoshi Ohnuki a,∗ , Hideaki Endo b , Mitsuru Izumi a a b

Faculty of Marine Technology, Tokyo University of Marine Science and Technology, 2-1-6 Etchujima, Koto-ku, Tokyo 135-8533, Japan Faculty of Marine Science, Tokyo University of Marine Science and Technology, 4-5-7 Konan, Minato-ku, Tokyo 108-8477, Japan

a r t i c l e

i n f o

Article history: Received 18 January 2012 Received in revised form 31 March 2012 Accepted 5 April 2012 Available online 12 April 2012 Keywords: Biosensor Glucose oxidase Self-assembled monolayer Langmuir–Blodgett film Prussian Blue

a b s t r a c t Ultrathin-form amperometric glucose biosensors have been fabricated on gold substrates by using selfassembled monolayers (SAMs) of carboxylic acid groups ( COOH). The glucose oxidase (GOx) was covalently immobilized on the SAMs and covered with Langmuir–Blodgett films including nanometersized clusters of Prussian Blue (PB) playing a role of mediator for glucose detection, which makes the present sensor, Au/SAM/GOx/PB, work at very low potentials of approximately 0.0 V vs. Ag/AgCl. An amperometric biosensor comprising a SAM of 4-mercaptobenzoic acid (MBA) exhibited a fast response current of 3 s, a detection limit of 12.5 ␮M, a high sensitivity of 50 nA/(cm2 mM), and a long-term stable linearity ranging from12.5 ␮M to 70 mM. The sensitivity is significantly affected by the conductivity of the SAM layer. The analysis of the electron transfer confirmed a high conductive nature of MBA among carboxylic acid groups. The obtained linearity in a high-concentration region is attributed to both the ultrathin nature and high electron transfer of the MBA SAM. © 2012 Elsevier B.V. All rights reserved.

1. Introduction In the past few decades, biosensors have received much attention due to their fast response, small size, and reliable and precise sensing in blood analyses [1]. The immobilization of enzymes like glucose oxidase (GOx) has been widely carried out in biosensors [2–6] and enzyme-based biofuel cells [7–9]. The method of immobilization determines sensor performance since electrical sensing signals pass through the binding part between the enzyme and the electrode. Many strategies, including electrostatic adsorption [10], covalent binding [11], electrochemical deposition [12,13], and the entrapment method [14], have been adopted to immobilize enzymes with highly active state. At present, diabetes represents a serious worldwide public health problem because it causes severe damage to the kidney, heart and vision. Thus, a glucose biosensor is very important, especially in clinical chemistry/medicine, where glucose concentration is used as a clinical indicator of diabetes [6]. By classical amperometric measurement, however, a high applied potential of ca. 0.6 V vs. Ag/AgCl is necessary to detect the electrochemical oxidation of H2 O2 produced by GOx. In this range, the current is easily disrupted by other chemical species, such as ascorbic acid, urea and

∗ Corresponding author. Tel.: +81 3 5245 7466; fax: +81 3 5245 7466. E-mail address: [email protected] (H. Ohnuki). 0925-4005/$ – see front matter © 2012 Elsevier B.V. All rights reserved. http://dx.doi.org/10.1016/j.snb.2012.04.018

other oxidative species [15–17]. To overcome this problem, the use of a variety of mediators has been proposed [18]. Prussian Blue (FeIII 4 [FeII (CN)6 ]3 , PB), an inorganic material, is a choice for its high sensitivity and selective detection for H2 O2 aside from its relatively low cost and ease of preparation [19]. Since Karyakin et al. [20] reported the development of an amperometric biosensor based on PB in 1994, PB has been widely applied in the field of bioelectrochemical analysis [19,21–24]. Ultrathin-form amperometric biosensors exhibit an advantage of a wide linear range of glucose detection [2]. The free diffusion of glucose and oxygen through the ultrathin film enhances the efficiency of GOx turnover, preventing a nonlinear response that follows the Michaelis–Menten equation [21,25]. For preparation purposes, the Langmuir–Blodgett (LB) technique produces ultrathin monomolecular organic films of amphiphilic molecules. We have successfully applied this technique in fabricating organic–inorganic hybrid films containing PB nanoscale clusters as mediators [10]. The as-fabricated LB films advantageously exhibit a large surface-to-volume ratio and a short travel distance for charge carriers to the electrode. Self-assembled monolayer (SAM) is another route to obtain monolayer films of biological molecules on substrates [26–28]. By selecting thiolates, the properties of SAMs can be easily controlled. Different types of amperometric biosensors including glucose, laccase and hydrogen peroxide sensors, have been prepared using SAMs formed by chemically bonding between the enzyme molecule and the electrode

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Fig. 1. (a) Structure of the biosensor with combined SAM and LB films. (b) 3-Mercaptopropionic acid (MPA), 11-mercaptoundecanoic acid (MUA) and 4-mercaptobenzoic acid (MBA).

[18]. Electrical signals pass through the SAMs and the electrical conducting property of the SAMs would significantly affect sensor performance. To our knowledge, the combination of LB films and SAM methods has seldom been approached in biosensor research [29,30]. In the present work, we fabricated ultrathin film form biosensors in which GOx was covalently immobilized on a substrate-bound SAM and further covered by LB films containing PB nanoclusters as mediators (Fig. 1(a)). The as-fabricated biosensors work at a low applied potential of 0.0 V (vs. Ag/AgCl) with a wide linear range of glucose concentrations and a highly stable sensing performance. To optimize the sensor performance, we examined three types of SAMs with different electrical conductivities: 11-mercaptoundecanoic acid (HS(CH2 )10 COOH, MUA), 3-mercaptopropionic acid (HS(CH2 )2 COOH, MPA) and 4-mercaptobenzoic acid (HS(C6 H4 )COOH, MBA). MPA and MUA differ in aliphatic chain length, while MBA differs from these two by the virtue of the aromatic ring (Fig. 1(b)). From our experiment results, it was found that the sensitivity of the biosensors remarkably changes when SAMs formed from molecules with different resistances are used.

2.2. Sample preparation

2. Experimental

2.2.3. Accumulation of PB-containing LB films A Teflon LB trough with an effective area of 516 cm2 (Filgen, Japan) was used for the LB film deposition of PB. Z-type LB films containing PB nanometer-sized clusters were obtained at a fixed surface pressure of 25 mN/m. After each transfer cycle, the substrate was dried in air for 20 min [10]. The sample obtained in this process was denoted as Au/SAM/GOx/PB.

2.1. Chemicals and solutions Glucose oxidase (GOx, E.C. 1.1.3.4 type VII, from Aspergillus niger), 3-mercaptopropionic acid (MPA), 11-mercaptoundecanoic acid (MUA), 4-mercaptobenzoic acid (MBA), N-(3-dimethylaminopropyl) ethylcarbodiimide hydrochloride (EDC), N-hydroxysuccinimide (NHS), octadecyltrimethylammonium bromide (ODTA), Prussian Blue (PB) and potassium ferrocyanides (II) and (III) were purchased from Sigma–Aldrich. The other chemicals used in this work were obtained from Wako Chem (Japan). Water was purified by using a Millipore water purification system. All the reagents used were of analytical grade and used as received without further purification. Phosphate buffer solutions (PBS) were prepared with 0.05 M KH2 PO4 and 0.1 M KCl, in which the pH was adjusted by adding 1 M KOH. Glass slides (1 cm × 2 cm, Corning Eagle 2000), on which an Au electrode (thickness 300 nm) was superposed on Cr (thickness 80 nm) pre-deposited by vacuum evaporation through a shadow mask, were used as substrates for the electrochemical study. The sensing area of the evaporated electrode was 0.28 cm2 . For infrared reflection–absorption spectroscopy (IR-RAS), larger glass slides (1 cm × 5 cm) were employed with the same thickness of Au.

2.2.1. Preparation of SAMs Before the modification of the working electrode, the Aupatterned substrates were ultrasonically washed for 2 min with a piranha solution composed of 1:3 (v/v) 30% H2 O2 /98% H2 SO4 , then rinsed with pure water and dried under a N2 gas stream. After being immersed in the SAM (SAM = MPA, MUA, or MBA) solution prepared in EtOH for 12 h, the Au substrate was rinsed with EtOH to remove the unreacted SAMs, then rinsed with pure water and dried under a N2 gas stream. Following the above process, we obtained different types of COOH-terminated SAM on Au, denoted as Au/SAM. 2.2.2. Immobilization of GOx The surface COOH groups of the SAM were activated by immersing the substrate in PBS (pH = 6.5) containing 30 mM EDC and 15 mM NHS for 1.5 h. Then, the substrate was immediately immersed in 5 mg/mL GOx solution (pH = 5.5 PBS) for 12 h at room temperature to attach GOx covalently to the SAM layer. Finally, the substrate was subsequently washed with PBS and pure water, and dried under a N2 stream. The sample at this stage was denoted as Au/SAM/GOx.

2.3. Measurement procedures Electrochemical measurements were performed using an ALS/CHI model 701B electrochemical analyzer. A conventional three-electrode system was used for each measurement: a Pt wire was employed as the auxiliary electrode, an Ag/AgCl electrode as the reference electrode, and Au/SAM/GOx/PB was employed as the working electrode. To characterize the SAM formation, cyclic voltammograms (CVs) of bare and modified Au electrodes were recorded in 100 mM KCl containing 2.5 mM K3 [Fe(CN)6 ] and K4 [Fe(CN)6 ] (1:1) as the redox probe or in 0.5 M H2 SO4 solution. For amperometric measurements, the working electrodes were held at a corresponding potential in PBS for about 30 min to achieve a steady background current.IR-RAS spectra were performed using a Nicolet 6700 FT-IR spectrometer equipped with an FT-80 RAS attachment. The spectra were recorded at a fixed incident angle

H. Wang et al. / Sensors and Actuators B 168 (2012) 249–255

251

-3

7x10 7

(a)

amide I

-3 14x1014

6

amide II

Au/MPA/GOx

4 3 2 1 0

1200

Absorbance (arb. unit)

Absorbance (arb. unit)

12

5

(b)

Au/MPA/GOx Au/MBA/GOx Au/MUA/GOx

10 8 6 4 2

COO 1400

C=O Au/MPA 1600

1800

2000

0

1200

-1

1400

1600

1800

2000

-1

Wave number (cm )

Wave number (cm )

Fig. 2. IR-RAS profiles of Au/MPA (a) and Au/SAM/GOx (SAM = MPA, MBA, and MUA) (b).

of 80◦ from the surface normal. All measurements were performed in a dry N2 atmosphere. Atomic force microscopy (AFM) which was employed to determine surface morphology was carried out in dynamic force mode under an air atmosphere using the Nano Search Microscope SFT3500SPM. The scanning rate was 0.3 Hz. AFM system software was used to analyze the image data. 3. Results and discussion

Fig. 2(b) shows the IR-RAS profiles of the Au/SAM/GOx (SAM = MPA, MBA, and MUA) samples. Amides I and II with comparative intensities were observed for each sample, which indicates that the amounts of immobilized GOx on these three samples were almost the same. As the CV experiments demonstrated, the percentage of defects or pinholes in the SAMs varies to a large extent depending on the type of SAM used. However, we deduced that the density of GOx immobilized does not depend on the type of SAM. In other words, the GOx density does not depend on the percentage of defects or pinholes in the SAMs.

3.1. Cyclic voltammograms 3.3. AFM The CVs of the monolayer-covered electrodes provide information about the presence of defects or pinholes in the SAMs [31,32]. By comparing the CV data (Supporting information) obtained in [Fe(CN)6 ]3−/4− (1:1) and 0.5 M H2 SO4 solutions, it was found that the number of defects or pinholes in the SAMs studied, namely, MUA, MBA, MPA, was markedly different. The order of surface coverage was found to be: MUA > MBA > MPA. This observation is qualitatively consistent with the results published in a previous report [32]. 3.2. IR-RAS Surface modification was investigated by IR-RAS. Fig. 2(a) shows the IR-RAS profiles of the samples of Au/MPA and Au/MPA/GOx. For Au/MPA, two characteristic absorption peaks of the carboxylic acid-terminated group were observed. The absorption peak at 1724 cm−1 was assigned as the C O stretch of the carboxylic acid group (COOH), while the peak observed at 1415 cm−1 was attributed to the absorption corresponding to a symmetric carboxylate (COO− ) stretch. The observed peaks indicate that some of the MPA molecules are in the deprotonated form [10,33,34]. After GOx was immobilized, two prominent peaks appeared at 1666 cm−1 and 1545 cm−1 in the IR-RAS spectrum of Au/MPA/GOx. These absorption peaks were attributed to the characteristic absorption of the amide bond: the peak at 1666 cm−1 corresponds to the C O stretching mode (Amide I) and the absorption at 1545 cm−1 corresponds to the N H bending mode (Amide II) [6]. The amide bonds are only attributable to the GOx molecule; the peak densities of the Amide I and II bands were thus used to evaluate the GOx quantity on the Au surface.

To confirm our deduction, we performed AFM measurements on the Au/SAM/GOx samples. Fig. 3 shows AFM images of the bare Au and Au/MPA/GOx surfaces, respectively. For the bare Au substrate, we observed that the surface was composed of contiguous grains with sizes of 100–150 nm. All the grains had an angular shape with flat planes and sharp edges without contamination; the surfaces of the plane were smooth and clean. In contrast to the bare Au image, the Au/MPA/GOx image shows many small particles attached on the grain surface. The particles were 20–30 nm in diameter and 5–6 nm in height. Since the surface of the MPA SAM (Au/MPA) gave the same AFM image as that of pure Au, these particles were assigned to GOx. It has been determined by

Fig. 3. AFM images of bare Au (a) and Au/MPA/GOx (b). The scanned area was 1000 nm × 1000 nm. The z-direction is expressed by the color shading, where bright denotes a higher feature and dark denotes a lower one.

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Au/MBA/GOx/PB Au/MPA/GOx/PB Au/MUA/GOx/PB

2

j (nA/cm )

-9 300 300x10

Fig. 4. Proposed surface morphology of Au/SAM/GOx. The amounts of immobilized GOx clusters are almost the same on the two surfaces: SAM with many defects (a) and closely packed SAM (b).

100

0

1200

1300

1400

1500

1600

1700

(a) 1800

Time (s)

Au/MBA/GOx/PB Au/MPA/GOx/PB Au/MUA/GOx/PB

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|j-j0| (nA/cm )

X-ray diffraction analysis that GOx from A. niger is a diametric protein with a molecular weight of 160 kDa and dimer dimensions of 7.0 nm × 5.0 nm × 8.0 nm [35]. Accordingly, the observed particles do not correspond to single molecules but GOx clusters in which several GOx molecules are linked with each other [36]. Based on the AFM profiles, we proposed the morphology of Au/SAM/GOx as shown in Fig. 4(a) and (b). The cluster has a very large surface area compared with the size of the SAM molecules; thus, it will sit over a large number of SAM molecules. In this situation, the defects and pinholes will not affect the amount of GOx immobilization, since several chemical bonds will be sufficiently strong to immobilize a single GOx cluster.

200

600

y =34.26 + 47.28x 2 R = 0.9994

400 2

y = 14.57 + 14.44x, R = 0.9913

200

3.4. Optimization of experimental variables

2

y = 0.58 + 3.47x, R = 0.9992

0 0 3000

2

3.5. Amperometric response of the biosensor Amperometric responses of the Au/SAM/GOx/PB biosensors were recorded by successively adding aliquots of a stock glucose solution (Fig. 5(a)). The amplitude of the response current density significantly varied according to the different SAMs used. Very small responses were observed with Au/MUA/GOx/PB with each glucose injection, while Au/MPA/GOx/PB showed a fair increased current density. Both samples reached a steady state within approximately 20 s. For Au/MBA/GOx/PB, a significant enhancement of current density and fast response current were observed within 3 s. The calibration curves of the Au/SAM/GOx/PB biosensors derived from the amperometric results are shown in Fig. 5(b). A stable linear relationship between the glucose concentration and current density was observed for each of the samples. As has already been shown in the amperometric profiles, the current density varies among different SAMs in the order of MUA < MPA < MBA. The detailed analyses showed that the ratio of MUA:MPA:MBA = 1.0:6.0:13.9. In the previous section, we confirmed that the amount of GOx immobilized on the different SAMs was constant. Thus, the results indicate that the sensitivity is directly affected by the SAM layer. 3.6. Effect of SAM layer on amperometric measurement The mechanism of redox cycles occurring at the electrode surface is shown in Fig. 6. Since the GOx amount is constant as shown in IR-RAS, the H2 O2 production rate is also constant for all the

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4 6 8 10 12 Glucose concentration (mM)

14

(b) 16

Au/MBA/GOx/PB S1 S2 S3

2500 |j-j0| (nA/cm )

In the previous sections, the effects of GOx immobilization onto different SAMs have been discussed. Other parameters, including the number of LB deposited PB films, and the detection of pH and applied potential, were investigated by amperometric measurements (Supporting information). The optimal conditions for glucose detection were found as follows: 5 layers of LB deposited PB films, pH = 7 PBS and an applied potential of 0.0 V (vs. Ag/AgCl) [17]. These conditions were used for the following measurements.

2000 1500 1000 500

(c) 0 0

20

40

60

80

Glucose concentration (mM) Fig. 5. (a) Amperometric responses of three biosensors to successive additions of stock glucose solution. Each glucose injection corresponds to a 1.0 mM glucose concentration increase for the test system. (b) Calibration curve of the Au/SAM/GOx/PB biosensors with successive injections of glucose solution measured in PBS (pH = 7) at 0.0 V vs. Ag/AgCl. (c) Wide calibration curves of the Au/MBA/GOx/PB biosensors (three samples: S1, S2, S3) with successive injections of glucose solution measured in PBS (pH = 7) at 0.0 V.

samples. The difference in current production among the three SAMs is due to the difference in reaction speed from H2 O2 to 2OH− occurring on the PB nanocluster surface. For this reductive reaction, an electron supply from the electrode surface is required; accordingly, the rate of electron transfer between the PB layer and the electrode surface, the SAM structure in our case, can directly affect the current flow rate. Note, that while the GOx immobilized on the SAM may block electron transfer, it was supposed that sufficient unoccupied space was left between the GOx clusters for electrons to pass through.

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and H2 O2 diffusion [2,40], these processes can be promoted by biosensors with a thin-film property. The total thickness of the fabricated biosensor Au/SAM/GOx/PB was calculated to be approximately 17 nm, in which the MBA layer was 0.6 nm, GOx was 5 nm and a single layer of the LB film deposited PB was 2.3 nm [10]. We consider that such linearity in a high-concentration region is attributed to both the ultrathin nature and good electron transfer property of the MBA SAM. On the other hand, the lowest sensitivity was observed with our biosensor (Table 1). We proposed that this low sensitivity was due to the limited amount of GOx immobilized on the hybrid system. 3.8. Reproducibility, stability and selectivity Fig. 6. Schematic representation of redox cycling at the electrode surface.

A calculation based on the tunnelling mechanism provides a set of ratios for different SAMs. By the coherent nonresonant tunnelling mechanism [37], the rate of electron transfer, kET , is described as:kET = k0 exp(− ˇd),where k0 , ˇ, and d are the preexponential factor, the decay constant of trans-conduction, and the thickness of the SAM layers, respectively. From the previous experiments performed using conducting AFM, the ˇ-values of benzylmercaptan and alkanethiol have been estimated to be 0.42 A˚ −1 [38] and 1.04 A˚ −1 [39], respectively, and the d values of MBA, MPA, and MUA ˚ 5.8 A, ˚ and 15.8 A, ˚ without a COOH group were estimated to be 6.0 A, respectively. From these parameters, we obtain the relative rates of electron transfer as MUA:MPA:MBA = 3.9 × 10−5 :1:43. This ratio shows that MUA forms a quite resistive SAM; on the other hand, MBA gives a conductive SAM. The difference in conductivity among the SAMs is a causal factor of the amperometric signals. A large ratio discrepancy from the theoretical calculations is suggested to be due to the large contact resistance between PB clusters and SAMs or to the high-density molecular packing of the SAMs, which is not considered in the calculation.

Sensing reproducibility was evaluated by injecting a l mM glucose solution into the test system. The relative standard deviations calculated from 11 repetitive measurements were 2.1%, 4.0% and 11.7% for the MBA-, MPA- and MUA-based sensors, respectively. To examine long-term stability, the response to 1 mM glucose was monitored for 17 days and the samples were kept in PBS (pH = 7) at 4 ◦ C when not in use. As shown in Fig. 7, all of the samples showed that the response current density was stable for at least 1 week. In particular, the sample with MBA remained at its original higher amplitude for 13–14 days. After this period, the output current intensity gradually decreased to 76.8% of its original signal on the 15th day, followed by a decrease to 44.2% on the 17th day. The selectivity was tested using ascorbic acid and uric acid (Supporting information). It was found that ascorbic acid produced an obvious interference current in the amperometric measurement. This is because our sample for studying the purpose of the thin-film was not covered with protective layer. For application of the Au/SAM/GOx/PB samples, further studies on the use of protective films [24,42] or feasibility of the routine analyses using real blood samples are required.

3.7. Advantage of MBA SAM for amperometric sensors

80

Au/MBA/GOx/PB Au/MPA/GOx/PB Au/MUA/GOx/PB

60 2

|j-j0| (nA/cm )

The above comparative studies demonstrate that the MBA SAM gives the highest sensitivity together with the shortest response time and the most stable performance (Fig. 7). However, only a few papers have been published employing MBA as a SAM for linking enzymes to a metal electrode [32]. Since it forms a COOHterminated SAM which can be used to establish a chemical bond with various proteins, MBA has considerable potential for sensor applications due to the aromatic moiety, C6 H4 , which enables the formation of MBA layers with a high conducting property. The profile of the Au/MBA/GOx/PB biosensor is defined by the linear regression equation I (nA/cm2 ) = 34.26 + 47.28 [glucose] (mM) with a correlation coefficient of 0.9994 (Fig. 5(b)). The detection limit was 12.5 ␮M (S/N = 3) with a sensitivity of 50 nA/(cm2 mM) and a stable linear relationship up to 70 mM (Fig. 5(c)). This sensing performance was compared with that previously reported for other thin-film structures (Table 1). Since the linearity is defined by the efficiency of GOx turnover, which depends on both O2 access

40 20 0 2

4

6

8 10 Days

12

14

16

Fig. 7. Stability of Au/SAM/GOx/PB layers (5) using response current densities to 1 mM glucose monitored in PBS (pH = 7) at 0.0 V.

Table 1 Performance with PB/GOx based thin-film biosensors. Film preparation

Detection limit (␮M)

Linear range (mM)

Sensitivity (nA/(cm2 mM))

Ref.

Sol–gel/PB/CNTa Layer-by-layerb LB films (4 layers)c SAM-LB

7.5 10 10 12.5

≤1.3 ≤11 ≤15 ≤70

1.52 × 104 3.2 × 103 1.2 × 102 5.0 × 10

[41] [21] [10] This work

a b c

Covalently cross-linked GOx–chitosan–SiO2 sol–gel composite film together with a PB/carbon nanotube hybrid layer. Layer-by-layer films of polymer protected PB and GOx. Electrostatically immobilized PB and GOx in ODTA LB films.

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4. Conclusions Preparation and performance of ultrathin-form amperometric glucose biosensors by the combined use of SAMs and LB techniques were studied. Among three SAM based biosensors fabricated using MPA, MUA and MBA, the best sensing performance was achieved with Au/MBA/GOx covered by 5 layers of LB deposited PB films in PBS (pH = 7) at a very low potential of 0.0 V vs. Ag/AgCl by the catalysis of PB. The sensitivity was supposed to be affected by the conductivity of the SAM layer. The best performance was observed with a fast response time of 3 s, a detection limit of 12.5 ␮M, a sensitivity of 50 nA/(cm2 mM), and a stable linearity of up to 70 mM, which was attributed to the ultrathin nature and high electron transfer caused by the existence of the aromatic ring of MBA. For wide application of SAMs in biosensing, the present work is expected to be a basis to further progress the development of amperometric biosensors based on SAMs. Acknowledgments This research was supported by a Grant-in-Aid for Scientific Research (B) (No. 2136006) from Japan Society for the Promotion of Science (JSPS) and by a Grant-in-Aid from the Faculty of Marine Technology, Tokyo University of Marine Science and Technology. Appendix A. Supplementary data Supplementary data associated with this article can be found, in the online version, at http://dx.doi.org/10.1016/j.snb.2012.04.018. References [1] P.R. Solanki, S.K. Arya, Y. Nishimura, M. Iwamoto, B.D. Malhotra, Cholesterol biosensor based on amino-undecanethiol self-assembled monolayer using surface plasmon resonance technique, Langmuir 23 (2007) 7398–7403. [2] J.J. Gooding, P. Erokhin, D.B. Hibbert, Parameters important in tuning the response of monolayer enzyme electrodes fabricated using self-assembled monolayers of alkanethiols, Biosensors and Bioelectronics 15 (2000) 229–239. [3] A.P.F. Turner, Biochemistry – biosensors sense and sensitivity, Science 290 (2000) 1315–1317. [4] Y. Liu, S. Wu, H. Ju, L. Xu, Amperometric glucose biosensing of gold nanoparticles and carbon nanotube multilayer membranes, Electroanalysis 19 (2007) 986–992. [5] E. Asav, E. Akyilmaz, Preparation and optimization of a bienzymic biosensor based on self-assembled monolayer modified gold electrode for alcohol and glucose detection, Biosensors and Bioelectronics 25 (2010) 1014–1018. [6] F.N. Comba, M.D. Rubianes, L. Cabrera, S. Gutierrez, P. Herrasti, G.A. Rivas, Highly sensitive and selective glucose biosensing at carbon paste electrodes modified with electrogenerated magnetite nanoparticles and glucose oxidase, Electroanalysis 22 (2010) 1566–1572. [7] D. Ivnitski, B. Branch, P. Atanassov, C. Apblett, Glucose oxidase anode for biofuel cell based on direct electron transfer, Electrochemistry Communications 8 (2006) 1204–1210. [8] I. Willner, Y.M. Yan, B. Willner, R. Tel-Vered, Integrated enzyme-based biofuel cells – a review, Fuel Cells 9 (2009) 7–24. [9] C. Tanne, G. Gobel, F. Lisdat, Development of a (PQQ)-GDH-anode based on MWCNT-modified gold and its application in a glucose/O2 -biofuel cell, Biosensors and Bioelectronics 26 (2010) 530–535. [10] H. Ohnuki, T. Saiki, A. Kusakari, H. Endo, M. Ichihara, M. Izumi, Incorporation of glucose oxidase into Langmuir–Blodgett films based on Prussian blue applied to amperometric glucose biosensor, Langmuir 23 (2007) 4675–4681. [11] C. Gouveia-Caridade, R. Pauliukaite, C.M.A. Brett, Development of electrochemical oxidase biosensors based on carbon nanotube-modified carbon film electrodes for glucose and ethanol, Electrochimica Acta 53 (2008) 6732–6739. [12] Z. Wang, S. Liu, P. Wu, C. Cai, Detection of glucose based on direct electron transfer reaction of glucose oxidase immobilized on highly ordered polyaniline nanotubes, Analytical Chemistry 81 (2009) 1638–1645. [13] M. Ammam, J. Fransaer, AC-electrophoretic deposition of glucose oxidase, Biosensors and Bioelectronics 25 (2009) 191–197. [14] M. Barbadillo, E. Casero, M.D. Petit-Domínguez, L. Vázquez, F. Pariente, E. Lorenzoa, Gold nanoparticles-induced enhancement of the analytical response of an electrochemical biosensor based on an organic–inorganic hybrid composite material, Talanta 80 (2009) 797–802. [15] B. Haghighi, S. Varma, F.M. Alizadeh Sh, Y. Yigzaw, L. Gorton, Prussian Blue modified glassy carbon electrodes – study on operational stability and its application as a sucrose biosensor, Talanta 64 (2004) 3–12.

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Biographies Huihui Wang studied organic chemistry at Shanghai University, Shanghai, China, where she received her master’s degree in 2008. In 2009, she joined Laboratory of Applied Physics, Tokyo University of Marine Science and Technology, Tokyo, Japan, as Ph.D. student. The main focus of her work is the fabrication and study of electrochemical biosensor based on the ultra-thin films. Hitoshi Ohnuki received his Ph.D. degree in science from University of Tsukuba, Ibaraki, Japan in 1998. He joined Tokyo University of Mercantile Marine, Tokyo, Japan from 1990 to 2003. He is currently an assistant professor in Tokyo University of Marine Science and Technology, Tokyo, Japan. His research interests is the development of variety of devices based on thin organic films including biosensors, organic field effect transistors, and other molecular based devices.

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Hideaki Endo received his Ph.D. from Tokyo Institute of Technology, Tokyo, Japan, in 1991. From 1991 he was an assistant professor at Tokyo University of Fisheries, Tokyo, Japan. From 1998 he was an associate professor at Tokyo University of Fisheries. Since 2011 he is a professor at Tokyo University of Marine Science and Technology, Tokyo, Japan. His current research interest is the development of biosensors for the field of marine and fisheries sciences. Izumi Mitsuru received his Ph.D. degree in solid-state physics from University of Tsukuba, Ibaraki, Japan, in 1983. He joined the Institute of Physics, University of Tsukuba in 1983. He moved to Nagasaki University, Nagasaki, Japan in 1986 and then to Tokyo University of Marine Science and Technology (TUMSAT), Tokyo, Japan in 1987. He is currently a professor and a director of Office of Liaison and Cooperative Research with TUMSAT. His research interests include a wide rage of topics in high-temperature superconductivity materials and its applications, and also the development of electrical devices based on thin organic films.