# Gold andTiO2 Nanostructurated Surfaces for Assembling of Electrochemical Biosensors

Authors: Antonella Curulli; Daniela Zane
Journal: Research Letters in Nanotechnology
 (2008)

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## Abstract

Devices based on nanomaterials are emerging as a powerful and general class of ultrasensitive sensors for the direct detection of biological and chemical species. In this work, we report the preparation and the full characterization of nanomaterials such as gold nanowires andTiO2 nanostructured films to be used for assembling of electrochemical biosensors. Gold nanowires were prepared by electroless deposition within the pores of polycarbonate particle track-etched membranes (PTMs). Glucose oxidase was deposited onto the nanowires using self-assembling monolayer as an anchor layer for the enzyme molecules. Finally, cyclic voltammetry was performed for different enzymes to test the applicability of gold nanowires as biosensors. Considering another interesting nanomaterial, the realization of functionalised TiO2 thin films on Si substrates for the immobilization of enzymes is reported. Glucose oxidase and horseradish peroxidase immobilized onto TiO2-based nanostructured surfaces exhibited a pair of well-defined and quasireversible voltammetric peaks. The electron exchange between the enzyme and the electrodes was greatly enhanced in the TiO2 nanostructured environment. The electrocatalytic activity of HRP and GOD embedded in TiO2 electrodes toward H2O2 and glucose, respectively, may have a potential perspective in the fabrication of third-generation biosensors based on direct electrochemistry of enzymes.


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## Body

## 1. Introduction

The use of nanoscale materials for electrochemical biosensors has seen explosive growth in
the past 5 years, since discovery of low-potential detection of NADH on carbon
nanotube-modified electrode by Musameh et al. [1] and the first use of gold nanoparticles
as labels for the electrochemical immunosensors by Dequaire et al.
[2]. Many research papers and general
reviews using nanomaterials for electrochemical biosensing [3, 4] have been
published since then.In this work, we report the characterization of nanomaterials such
as gold nanowires and TiO2 nanomaterials to be used for assembling
of electrochemical biosensors and the comparison between the two systems used.
The immobilization of single or clustered molecules of enzymes on the
nanostructured surfaces could be an important step towards the development of a
new class of biosensors.


## 2. Experimental

Further details of the experimental
part have been mentioned in the Supplementary
Material (available online at doi:10.1155/2008/789153).The chemical analysis (X-ray
photoelectron spectroscopy) was performed using an Escalab Mk II spectrometer,
equipped with a standard Al Kα excitation source (hν=1486.6 eV) and a five-channeltron detection system.The FT-IR spectra were recorded with an IR Scope II Bruker,
Interferometer Equinox 55 (Bruker).The electrochemical measurements were performed at room temperature
in a conventional one-compartment cell with a three-electrode configuration
using an Autolab PGSTAT 12 potentiosta/galvanostat (Eco Chemie, the Netherlands).
The reference electrode was Ag/AgCl and a Pt electrode was the counter.


## 3. Results and Discussion

### 3.1. Gold Nanowires

The synthesis and morphological characterization
of gold nanowires have been discussed elsewhere [5]. In this work, we focused
on nanowires grown with a synthesis time of 24 hours.After the preparation of the nanowires film, the next steps towards
the biosensor fabrication are the SAM formation and subsequent activation to
allow the attachment of the enzyme. The resulted nanowires system should allow
the diffusion and the immobilization of the GOD, a large enzyme [6]. Hou et
al. [7] demonstrated that the films prepared by electroless deposition can be
used as substrates to support densely packed SAMs formed from long chains. In
this paper, short-chain alkane thiols have been used because the long-chain
systems have the drawback of a great distance of the enzyme redox centre from
the electrode surface and of a limited diffusion.3-mercaptopropionic acid (MPA) and 2-mercaptoethylamine
(MPE) were used to link the enzyme onto the surface.
The formation of SAMs on electroless has been qualitatively
assessed by measurements of contact angles of water. For
bare electroless gold, a contact angle
of 75 ± 8° was found.
This value decreases to 32 ± 3° for Nano-Au/MPE and to less than 10° for Nano-Au/MPA.
This modification of the
contact angles is, therefore, a useful proof of the
presence of hydrophilic at the surface of electroless
gold [8].The effective surface area of electrochemically active gold
supporting SAM was estimated by cyclic voltammetry in 0.1 M H2SO4.
The Nano-Au/MPE system has a fractional coverage of 74% and the Nano-Au/MPA has
a coverage of 55% according to the literature data [9].In Tables1 and 2, the values of the peak current for nanostructured
and non-nanostructured systems are reported, named, respectively, Ipa and
Io. We observed a ratio between the peak current in the presence and
in the absence of the nanostructured film ranging from 9.79 up to 15.23, which indicates an increase of
the elctrochemical response using the nanowired system. Similar data were
obtained for the other systems, not shown here for brevity.Table 1
Electrochemical  characterization of gold
nanowires-modified electrodes with SAM (MPA in KCl 0.2 M).Ipa is the
peak current at gold nanowires SAM-modified electrode, Io is the
peak current for non-nanostructured electrode.
Epa (V)Ipa (μA)Ipa/IoAu/MPANano-Au/MPAAu/MPANano-Au/MPAFe(CN)63−0.270.3613.83210.6415.23Ru(NH3)63+−0.11−0.3320.42200.009.79Table 2
Electrochemical
characterization of gold nanowires-modified electrodes with SAM (MPE in KCl 0.2 M).Ipa is the
peak current at gold nanowires SAM-modified electrode,
Io is the
peak current for non-nanostructured electrode.
Epa (V)Ipa (μA)Ipa/IoAu/MPENano-Au/MPEAu/MPENano-Au/MPEFe(CN)63−0.280.3518.14187.7610.35Ru(NH3)63+−0.10−0.0421.70252.1311.62Enzyme immobilization and
electrochemical detection of glucose
In the system modified with −NH2 terminal thiol
(Nano-Au/MPE), the enzyme molecules were covalently attached using
glutaraldehyde as linking agent, while the covalent attachment of GOD to the carboxylic-terminated
SAM was achieved via carbodiimide activation (EDC + NHS). In this case, the
physical adsorption immobilization failed the stability test during time: in
fact, in our experimental conditions, the biosensors obtained have worked only
for few hours. In any case, generally, the physical adsorption has not been
used as enzyme-immobilizing system for SAM-modified surface [8, 9].The different enzyme electrodes have
then been used to detect the target analyte (glucose) by cyclic voltammetry
ranging the analyte concentrations from 1 to 6 mM.The highest current response was
seen for the Nano-Au/MPE/GOD system and the corresponding regression linear
equation is the following:y=68.629x+28.61,
where y represents the current in μA and x the
glucose concentration in mM, R2 was 0.9987. This high
response could be due to the higher roughness of the nanostructured surface
compared to a flat gold electrode and to the tubular nature of the electrode,
both leading to an increased surface area of the electrode for the same
geometric area [10]. Another possible explanation of the high sensitivity
observed with our system could be that the amount of the active enzyme
molecules is higher when they are immobilized within porous system than on a
flat surface. The difference observed between the two immobilized systems is
probably due to the fact that the glutaraldehyde system probably loads a lower
quantity of the enzyme because the linker agent is largely diluted before being
used.


### 3.2. TiO2-Based Nanomaterials

The morphological characterization
of TiO2-modified electrodes was discussed elsewhere [11, 12].The enzymes have been immobilized by
physical adsorption because the functionalised surface should be an optimal
platform for anchoring enzymes. The effectiveness of enzyme immobilization was
determined by XPS and FT-IR measurements, performed on TiO2-GOD- and TiO2-HRP-modified electrodes.
From XPS data, three components (C1, C2, and C3) were observed, corresponding,
respectively, to the aliphatic bond CH–CH present in GOD
structure, and to the different carbon atoms of the characteristic amidic
groups [13]. The presence of adsorbed GOD enzyme on TiO2-modified Si
electrode was confirmed by FT-IR spectra, showing the characteristic amide I
band (1649 cm−1) due to the stretching vibration of peptide C=O groups,
and amide II band (1556 cm−1), which results from a combination of N–H in-plane
bending and C–N stretching of
peptide groups [14]. We observed only a slight shift with respect to the amide
peak positions of GOD powder so that the immobilization process used retains
the native GOD conformation [15], but we could suggest a strong interaction
between the enzyme and the surface just to justify the shift of the amide peak
position. Similar results (not shown for the sake of brevity) were obtained for
the immobilized HRP enzyme.As shown in Figure1, direct
electron transfer of GOD and HRP at TiO2-modified Si electrodes was
observed. The redox peaks potentials are consistent with reported values for
free FAD and FAD redox centre for flavoenzyme [16]. The peak positions remain
constant as the scan rate increases, while the redox peak currents are proportional
to the scan rate in the range less than 100 mV/s, indicating a typical
surface-controlled quasireversible process [17].Cyclic voltammograms for (a) GOD-TiO2-modified
Si electrodes, (b) HRP-TiO2-modified Si electrodes at different scan
rate and at pH 7.0 in 0.1 M phosphate buffer solution.
(a)(b)The amperometric response of GOD and
HRP third-generation biosensor based on functionalised TiO2–Si electrodes are
shown in Figure 2. The calibration curve for glucose (Figure 2(a)) and for H2O2 (Figure 2(b)) showed a linear response for the biosensor in the concentration
range of 5.0×10−6 to 5.5×10−4 M for glucose and 1.0×10−6 to 2.0×10−4 M
for HRP, with a time response of few seconds (7 seconds for GOD and 6 seconds
for HRP) for successive enzyme injection. At higher glucose and H2O2 concentrations, the calibration curves tend to level off, showing a typical Michaelis-Menten mechanism.
The apparent Michaelis-Menten constant
(Km) was calculated at different enzyme concentration. Small Km values were
obtained for 5 mg/mL of GOD (Km = 7.5 mM) and HRP (Km = 1.0 mM) concentrations,
indicating a great affinity between enzyme and electrode surface. Decreasing the
enzyme concentration, Km values increase up to, respectively, 30 mM and 5.8 mM for 1 mg/mL of GOD
and HRP concentration, indicating that glucose and H2O2 biosensors still work well at low enzyme concentration.Calibration curve of (a) glucose biosensor based on
TiO2-modified Si electrodes, and (b) H2O2 biosensor based on TiO2-modified Si electrodes. Inset graphs show
typical amperometric response of the biosensors to successive injection of
glucose and H2O2 into stirring 0.1 M phosphate buffer
solution at pH 7.0.
(a)(b)


## 3.1. Gold Nanowires

The synthesis and morphological characterization
of gold nanowires have been discussed elsewhere [5]. In this work, we focused
on nanowires grown with a synthesis time of 24 hours.After the preparation of the nanowires film, the next steps towards
the biosensor fabrication are the SAM formation and subsequent activation to
allow the attachment of the enzyme. The resulted nanowires system should allow
the diffusion and the immobilization of the GOD, a large enzyme [6]. Hou et
al. [7] demonstrated that the films prepared by electroless deposition can be
used as substrates to support densely packed SAMs formed from long chains. In
this paper, short-chain alkane thiols have been used because the long-chain
systems have the drawback of a great distance of the enzyme redox centre from
the electrode surface and of a limited diffusion.3-mercaptopropionic acid (MPA) and 2-mercaptoethylamine
(MPE) were used to link the enzyme onto the surface.
The formation of SAMs on electroless has been qualitatively
assessed by measurements of contact angles of water. For
bare electroless gold, a contact angle
of 75 ± 8° was found.
This value decreases to 32 ± 3° for Nano-Au/MPE and to less than 10° for Nano-Au/MPA.
This modification of the
contact angles is, therefore, a useful proof of the
presence of hydrophilic at the surface of electroless
gold [8].The effective surface area of electrochemically active gold
supporting SAM was estimated by cyclic voltammetry in 0.1 M H2SO4.
The Nano-Au/MPE system has a fractional coverage of 74% and the Nano-Au/MPA has
a coverage of 55% according to the literature data [9].In Tables1 and 2, the values of the peak current for nanostructured
and non-nanostructured systems are reported, named, respectively, Ipa and
Io. We observed a ratio between the peak current in the presence and
in the absence of the nanostructured film ranging from 9.79 up to 15.23, which indicates an increase of
the elctrochemical response using the nanowired system. Similar data were
obtained for the other systems, not shown here for brevity.Table 1
Electrochemical  characterization of gold
nanowires-modified electrodes with SAM (MPA in KCl 0.2 M).Ipa is the
peak current at gold nanowires SAM-modified electrode, Io is the
peak current for non-nanostructured electrode.
Epa (V)Ipa (μA)Ipa/IoAu/MPANano-Au/MPAAu/MPANano-Au/MPAFe(CN)63−0.270.3613.83210.6415.23Ru(NH3)63+−0.11−0.3320.42200.009.79Table 2
Electrochemical
characterization of gold nanowires-modified electrodes with SAM (MPE in KCl 0.2 M).Ipa is the
peak current at gold nanowires SAM-modified electrode,
Io is the
peak current for non-nanostructured electrode.
Epa (V)Ipa (μA)Ipa/IoAu/MPENano-Au/MPEAu/MPENano-Au/MPEFe(CN)63−0.280.3518.14187.7610.35Ru(NH3)63+−0.10−0.0421.70252.1311.62Enzyme immobilization and
electrochemical detection of glucose
In the system modified with −NH2 terminal thiol
(Nano-Au/MPE), the enzyme molecules were covalently attached using
glutaraldehyde as linking agent, while the covalent attachment of GOD to the carboxylic-terminated
SAM was achieved via carbodiimide activation (EDC + NHS). In this case, the
physical adsorption immobilization failed the stability test during time: in
fact, in our experimental conditions, the biosensors obtained have worked only
for few hours. In any case, generally, the physical adsorption has not been
used as enzyme-immobilizing system for SAM-modified surface [8, 9].The different enzyme electrodes have
then been used to detect the target analyte (glucose) by cyclic voltammetry
ranging the analyte concentrations from 1 to 6 mM.The highest current response was
seen for the Nano-Au/MPE/GOD system and the corresponding regression linear
equation is the following:y=68.629x+28.61,
where y represents the current in μA and x the
glucose concentration in mM, R2 was 0.9987. This high
response could be due to the higher roughness of the nanostructured surface
compared to a flat gold electrode and to the tubular nature of the electrode,
both leading to an increased surface area of the electrode for the same
geometric area [10]. Another possible explanation of the high sensitivity
observed with our system could be that the amount of the active enzyme
molecules is higher when they are immobilized within porous system than on a
flat surface. The difference observed between the two immobilized systems is
probably due to the fact that the glutaraldehyde system probably loads a lower
quantity of the enzyme because the linker agent is largely diluted before being
used.


## 3.2. TiO2-Based Nanomaterials

The morphological characterization
of TiO2-modified electrodes was discussed elsewhere [11, 12].The enzymes have been immobilized by
physical adsorption because the functionalised surface should be an optimal
platform for anchoring enzymes. The effectiveness of enzyme immobilization was
determined by XPS and FT-IR measurements, performed on TiO2-GOD- and TiO2-HRP-modified electrodes.
From XPS data, three components (C1, C2, and C3) were observed, corresponding,
respectively, to the aliphatic bond CH–CH present in GOD
structure, and to the different carbon atoms of the characteristic amidic
groups [13]. The presence of adsorbed GOD enzyme on TiO2-modified Si
electrode was confirmed by FT-IR spectra, showing the characteristic amide I
band (1649 cm−1) due to the stretching vibration of peptide C=O groups,
and amide II band (1556 cm−1), which results from a combination of N–H in-plane
bending and C–N stretching of
peptide groups [14]. We observed only a slight shift with respect to the amide
peak positions of GOD powder so that the immobilization process used retains
the native GOD conformation [15], but we could suggest a strong interaction
between the enzyme and the surface just to justify the shift of the amide peak
position. Similar results (not shown for the sake of brevity) were obtained for
the immobilized HRP enzyme.As shown in Figure1, direct
electron transfer of GOD and HRP at TiO2-modified Si electrodes was
observed. The redox peaks potentials are consistent with reported values for
free FAD and FAD redox centre for flavoenzyme [16]. The peak positions remain
constant as the scan rate increases, while the redox peak currents are proportional
to the scan rate in the range less than 100 mV/s, indicating a typical
surface-controlled quasireversible process [17].Cyclic voltammograms for (a) GOD-TiO2-modified
Si electrodes, (b) HRP-TiO2-modified Si electrodes at different scan
rate and at pH 7.0 in 0.1 M phosphate buffer solution.
(a)(b)The amperometric response of GOD and
HRP third-generation biosensor based on functionalised TiO2–Si electrodes are
shown in Figure 2. The calibration curve for glucose (Figure 2(a)) and for H2O2 (Figure 2(b)) showed a linear response for the biosensor in the concentration
range of 5.0×10−6 to 5.5×10−4 M for glucose and 1.0×10−6 to 2.0×10−4 M
for HRP, with a time response of few seconds (7 seconds for GOD and 6 seconds
for HRP) for successive enzyme injection. At higher glucose and H2O2 concentrations, the calibration curves tend to level off, showing a typical Michaelis-Menten mechanism.
The apparent Michaelis-Menten constant
(Km) was calculated at different enzyme concentration. Small Km values were
obtained for 5 mg/mL of GOD (Km = 7.5 mM) and HRP (Km = 1.0 mM) concentrations,
indicating a great affinity between enzyme and electrode surface. Decreasing the
enzyme concentration, Km values increase up to, respectively, 30 mM and 5.8 mM for 1 mg/mL of GOD
and HRP concentration, indicating that glucose and H2O2 biosensors still work well at low enzyme concentration.Calibration curve of (a) glucose biosensor based on
TiO2-modified Si electrodes, and (b) H2O2 biosensor based on TiO2-modified Si electrodes. Inset graphs show
typical amperometric response of the biosensors to successive injection of
glucose and H2O2 into stirring 0.1 M phosphate buffer
solution at pH 7.0.
(a)(b)


## 4. Conclusions

Glucose oxidase was immobilized on the nanostructured gold-modified
surfaces via self-assembling monolayers. Electrochemical analysis showed that a
great enhancement of the peak current occur if the sensitive material is
deposited on a nanostructured surface, which makes gold nanowired surfaces
excellent candidates for biosensing applications, but further work should be
necessary to improve the stability of our system.On the other hand, nanostructured
TiO2-based layers provide an efficient immobilization of HRP and GOD
monolayers on Si platforms, without the use of any mediators, considering the
literature data for other glucose electrochemical sensors [18], the obtained
results resulted very promising. The direct electron transfer between enzyme
and electrodes is enhanced by the nanostructured environment of TiO2-based
layers. Amperometric glucose and H2O2 third-generation
biosensors based on the TiO2-modified Si-electrodes showed good
sensitivity, low detection limit (10−6 M), and fast time response (few seconds).



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*Source: 789153-2008-04-03.xml*