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Electrochemical Immunochip Sensor for AflatoxinM1 Detection

Charlie O. Parker, Yvonne H. Lanyon, Mary Manning, Damien W. M. Arrigan,*, and

Ibtisam E. Tothill*,

Cranfield Health, Cranfield University, Cranfield, Bedfordshire, MK43 0AL, U.K., and Tyndall National Institute,Lee Maltings, University College Cork, Cork, Ireland

An investigation into the fabrication, electrochemical

characterization, and development of a microelectrode

array (MEA) immunosensor for aflatoxin M1 is presented

in this paper. Gold MEAs (consisting of 35 mi-

crosquare electrodes with 20 m 20 m dimensions

and edge-to-edge spacing of 200m) together with on-

chip reference and counter electrodes were fabricated

using standard photolithographic methods. The MEAs

were then characterized by cyclic voltammetry, and thebehavior of the on-chip electrodes were evaluated. The

microarray sensors were assessed for their applicabil-

ity to the development of an immunosensor for the

analysis of aflatoxin M1 directly in milk samples.

Following the sensor surface silanization, antibodies

were immobilized by cross-linking with 1,4-phenylene

diisothiocyanate (PDITC). Surface characterization was

conducted by electrochemistry, fluorescence micros-

copy, scanning electron microscopy (SEM), and atomic

force microscopy (AFM). A competitive enzyme linked

immunosorbent assay (ELISA) assay format was de-

veloped on the microarray electrode surface using the3,3,5,5-tetramethylbenzidine dihyrochloride (TMB)/

H2O2 electrochemical detection scheme with horserad-

ish peroxidase (HRP) as the enzyme label. The per-

formance of the assay and the microarray sensor were

characterized in pure buffer conditions before applying

to the milk samples. With the use of this approach,

the detection limit for aflatoxin M1 in milk was esti-

mated to be 8 ng L-1, with a dynamic detection

range of 10-100 ng L-1, which meets present legisla-

tive limits of 50 ng L-1. The milk interference with the

sensor surface was also found to be minimal. These

devices show high potential for development of a rangeof new applications which have previously only been

detected using elaborate instrumentation.

Electrochemical sensors are renowned for their excellent

sensitivity, selectivity, versatility, and simplicity, and therefore

there is a continual interest in their development for the analysis

of environmental, food, and clinical samples.1-6 Different types

of sensor platforms have been used for electrochemical sensors,

but most are based on screen-printed technology.7 Interest in the

use of microelectrodes based on photolithographic techniques

coupled with electrochemical detection methods is increasing.8-10

A microelectrode is described as an electrode where one of its

dimensions is in the micrometer range.11 These developments

and advances in sensor technology have been fuelled by medical

applications where microelectrodes can be implanted to monitorelectrophysiological pulses such as in cardiac tissues and also in

other applications.12 One of the main benefits of using a micro-

electrode in a sensor application is the greater sensitivity that

arises from the enhanced mass-transport at these small elec-

trodes.9 Hemispherical diffusion layers are formed at such

electrodes and a much faster diffusion of electroactive substances

occurs due to the multidimensional nature of this process,

resulting in sigmoidal (or steady-state) cyclic voltammograms

(CVs).13,14 The advantages are in the improved response time

(faster response), greater sensitivity and increased response per

unit electrode surface area (greater current density, increasing

the signal-to-noise ratio). However, this results in very low currentvalues which can be problematic.11,14 The use of an array of

microelectrodes addresses this problem by providing a substantial

improvement in the signal-to-noise ratio under steady-state

conditions.15,16The spacing between the electrodes in these arrays

* To whom correspondence should be addressed. Ibtisam E. Tothill: fax

+44(0)1234 75 8380, e-mail [email protected]. Damien W. M. Arrigan: fax

+353-21-4270271, e-mail [email protected]. Cranfield University. Tyndall National Institute.

(1) Tothill, I. E., Piletsky, S., Magan, N., Turner, A. P. F. In Instrumentation

and Sensors for the Food Industry, 2nd ed.;Woodhead Publishing Limited

CRC Press: Boca Raton, FL, 2001; pp 760-775.

(2) Mascini, M. Pure Appl. Chem. 2001, 73, 2330.(3) Wang, J. Acc. Chem. Res. 2002, 811816.(4) Tothill, I. E., Turner, A. P. F. In Encyclopedia of Food Sciences and Nutrition,

2nd ed.; Academic Press: New York, 2003; pp 489-499.

(5) Pemberton, R. M.; Mottram, T. T.; Hart, J. P. J. Biochem. Biophys. Methods

2005, 63, 201212.(6) Bakker, E.; Qin, Y. Anal. Chem. 2006, 78, 39653983.(7) Newman, J. D.; Turner, A. P. F. Biosens. Bioelectron. 2005, 20, 24352453.(8) Berduque, A.; Lanyon, Y. H.; Beni, V.; Herzog, G.; Watson, Y. E.; Rodgers,

K.; Stam, F.; Alderman, J.; Arrigan, D. W. M. Talanta 2007, 71, 1022

1030.(9) Ordeig, O.; del Campo, J.; Munoz, F. X.; Banks, C. E.; Compton, R. G.

Electroanalysis 2007, 19, 19731986.(10) Beni, V.; Arrigan, D. W. M. Current Anal. Chem. 2008, 4, 229241.(11) Stulik, K.; Amatore, C.; Holub, K.; Marecek, V.; Kutner, W. Pure Appl. Chem.

2000, 72, 14831492.(12) Hoffman, B. F. Cardiovasc. Res. 2002, 53, 15.(13) Amatore, C. In Physical Electrochemistry: Principles, Methods and Applica-

tions; Rubinstein, I., Ed.; Marcel Dekker: New York, 1995; p 131.

(14) Alden, J. A.; Booth, J.; Compton, R. G.; Dryfe, R. A. W.; Sanders, G. H. W.

J. Electroanal. Chem. 1995, 389, 4554.(15) Feeney, R.; Kounaves, S. P. Electroanalysis 2000, 12, 677684.

Anal. Chem. 2009, 81, 52915298

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is important and needs to be such that each element of the array

experiences individual, noninteracting diffusion profiles.16-19Then

steady-state behavior can be achieved, which is best for sensor

applications.

The enhanced capability of microelectrode arrays (MEAs) as

sensing devices makes them an ideal choice for trace analysis,

such as aflatoxin M1 (AFM1) analysis. Aflatoxin M1 (Cyclopenta

(C) furo(3,2:4,5)furo(2,3-H)(1)benzopyran-1,11-dione,2,3,6A,9A

tetrahydro-9-hydroxy-4-methoxy, CAS number 6795-23-9, chemi-

cal formula C17H12O7, relative molecular mass 328.3 amu) isexcreted in milk by animals upon the digestion of feed

contaminated with the fungal toxin aflatoxin B1.20,21 It has been

theorized that aflatoxin M1 is a detoxification product of aflatoxin

B1 since the carcinogenicity of aflatoxin M1 is lower than

aflatoxin B1.22 However, aflatoxin M1 is still regarded as a

carcinogenic, genotoxic, teratogenic, and immunosuppressive

compound. Aflatoxin M1 can also be found in other dairy

products such as cheese, yogurt, and infant formulas23 and also

in human breast milk.24 Because of the fact that the milk intake

in infants is high and that they are very vulnerable to toxins,

the European Commission regulation 472/2002 imposes maxi-

mum permissible levels of aflatoxin M1 in milk of 50 ng L

-1

and in infant formulas of 25 ng L-1.25

Determination of aflatoxin M1 is usually conducted using

HPLC, TLC, and ELISA methods which are all laboratory-based

systems and require the expertise of trained personnel.26-28

Unfortunately the regions of the world which are most affected

by aflatoxin contamination tend to be poorer areas within the

tropics. Therefore, as stipulated by the United Nations there is

an urgent need for simple, robust, low-cost analysis methods, for

the major mycotoxins, which can be used in developing country

laboratories.29 Microfabricated sensor systems offer many ben-

efits to achieve those goals.30

In this article, the development of an MEA-based immunosen-

sor for aflatoxin M1 is reported. The chip-based electrochemical

cell was fabricated to contain the working electrode, which was

the MEA, and counter and reference electrodes, so that all

necessary electrodes for electrochemical measurements were

contained on the chip. The assay on the sensor chip was based

on a competitive format between the free aflatoxin M1 in the

sample and an aflatoxin-horseradish peroxidase conjugate for

an immobilized monoclonal antibody for aflatoxin M1. With

the use of chronoamperometry, the depletion of hydrogen

peroxide was monitored via 3,3,5,5-tetramethylbenzidine di-

hyrochloride (TMB) mediation to ascertain the concentrationof HRP on the sensor and consequently the concentration of

aflatoxin M1 in the sample.

EXPERIMENTAL SECTION

Reagents and Solutions. Aflatoxin M1 was purchased from

Axxora UK Limited (Nottingham, U.K.), anti-aflatoxin M1antibody (raised from rat) was purchased from Abcam Limited,

(Cambridge U.K.), and aflatoxin M1-HRP conjugate was

obtained from a RIDASCREEN kit from R-Biopharm (Glasgow,

U.K.). 3,3,5,5-Tetramethylbenzidine dihydrochloride, hydrogen

peroxide, and Tween 20 were purchased from Sigma-Aldrich

(Poole, U.K.). Anti-rat immunopure antibody (raised in goat

with affinity for the Fc fragment only) was from Perbio Science

(Cramlington, U.K.). Milk and dried milk samples were

obtained from the local supermarket. All other chemicals were

purchased from Sigma-Aldrich (Poole, U.K.) or otherwise as

stated in the text.

Microfabrication. Gold cell-on-a-chip microelectrodes (includ-

ing on-chip reference and counter electrodes) were fabricated by

standard deposition, etching, and lithographic techniques used

in microfabrication technology. The first step involved growth of

a thermal oxide on a silicon wafer. This was followed by plasma

enhanced chemical vapor deposition (PECVD) of a silicon dioxide

layer. Photoresist was then spun onto the wafer and patterned,

and the exposed oxide layer was then wet etched. For thefabrication of the metal electrodes, gold was deposited by

evaporation of Ti/Pt/Au multilayers in the proportion 30:50:250

nm and the remaining silicon dioxide was removed by a buffer

oxide etch (HF and NH4F). This was followed by deposition of

a Si3N4 passivation layer. The recessed microelectrode array

(500 nm recess depth) was then obtained using a photolitho-

graphic etch process (Pt-EKC solvent). Following fabrication,

the wafers were diced and the electrodes packaged on printed

circuit boards (PCB) by attaching the individual chips to the

PCB with silver epoxy die attach (Ablebond 8484, Ablestik),

wire bonding the bondpads on the chips to the PCB with 25

m aluminum wires, and finally protecting the wirebonds andchip edges by covering in a polymeric selective encapsulant

(Amicon 50300 HT, Emerson & Cuming). The electrode device

containing all necessary electrodes (working, counter, and

reference) is referred to as the cell-on-chip device. The

electrode materials were either gold or platinum, including the

pseudoreference electrodes. All microfabrication processing

was carried out at the Central Fabrication Facility at Tyndall

National Institute (Cork, Ireland).

(16) Bard, A. J., Faulkner, L. R. In Fundamentals and Applications, 2nd ed.; John

Wiley and Sons: New York, 2001; p 168.

(17) Sandison, M.; Anicet, N.; Glidle, A.; Cooper, J. M. Anal. Chem. 2002, 74,

57175725.(18) Davies, T. J.; Compton, R. G. J. Electroanal. Chem. 2005, 585, 6382.(19) Davies, T. J.; Ward-Jones, S.; Banks, C. E.; del Campo, J.; Mas, R.; Munoz,

F. X.; Compton, R. G. J. Electroanal. Chem. 2005, 585, 5162.(20) Sargeant, K.; Sheridan, A.; OKelly, J. Nature 1961, 192, 10961097.(21) Holzapfel, C. W.; Steyn, P. S. Tetrahedron Lett. 1966, 25, 27992803.(22) Neal, G. E.; Eaton, D. L.; Judah, D. J.; Verma, A. Toxicol. Appl. Pharmacol.

1998, 151, 152158.(23) Martins, M. L.; Martins, H. M. Int. J. Food Microbiol. 2004, 91, 315317.(24) El-Nezam, H. S.; Nicoletti, G.; Neal, G. E.; Donohue, D. C.; Ahokas, J. T.

Food Chem. Toxicol. 1995, 33, 173179.(25) Henry, S. H.; Whitaker, T.; Rabbani, I.; Bowers, J.; Park, D.; Price, W.; Bosch,

F. X.; Pennington, J.; Verger, P.; Yoshizawa, T.; van Egmond, H.; Jonker,

M. A.; Coker, R. Aflatoxin M1; Report 1012, (WHO Additives, Series 47),

Joint Expert Committee on Food Additives (JECFA), 2001.

(26) Kamkar, A. Food Control 2005, 16, 593599.(27) Oveisi, M. R.; Jannat, B.; Sadeghi, N.; Hajimahmoodi, M.; Nikzad, A. Food

Control2007, 18, 12161218.(28) Rodriguez Velasco, M. L.; Calonge Delso, M. M.; Ordonez Escudero, D.

Food Addit. Contam. 2003, 20, 276280.(29) Proctor, D. L., Ed. Grain Storage Techniques: Evolution and Trends in

Developing Countries; Food and Agriculture Organization of the United

Nations: Rome, Italy, 1994.

(30) Logrieco, A.; Arrigan, D. W. M.; Brengel-Pesce, K.; Siciliano, P.; Tothill,

I. E. Food Addit. Contam. 2005, 22, 335344.(31) Lucarelli, F.; Marrazza, G.; Turner, A. P. F.; Mascini, M. Biosens Bioelectron.

2004, 19, 515530.

(32) Yang, M.; Yau, H. C. M.; Chan, H. L. Langmuir1998, 14, 61216129.(33) Ouerghi, O.; Touhami, A.; Othmane, A.; Ben Ouada, H.; Martelet, C.;

Fretigny, C.; Jaffreezic Renault, N. Biomol. Eng. 2002, 19, 183188.

(34) Parker, C.; Tothill, I. E. Biosensors. Bioelectron. 2009, 24, 24522457.

5292 Analytical Chemistry, Vol. 81, No. 13, July 1, 2009

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Electrochemical Characterization and Reagents. A CHI620A

electrochemical analyzer with picoamp booster and faraday cage

(CH Instruments, Texas) was used for the electrochemical

characterization studies. Experiments were performed either with

external reference and counter electrodes with a Ag|AgCl refer-

ence electrode and platinum wire counter electrode (both from

CH instruments) or using the on-chip (pseudo)reference and

counter electrodes. Both on-chip pseudoreference and off-chip

reference electrodes were used in the evaluation study so as to

provide a realistic comparison to standard laboratory operations.Prior to electrochemical testing, the chips were treated with

oxygen plasma for 10 min at 150 W to remove any residual organic

matter. Cyclic voltammograms (CVs) were recorded at the Au

MEAs in 1 mM ferrocene monocarboxylic acid (FcCOOH) in 0.01

M phosphate buffered saline (PBS) solution. Characterization of

DNA-surface modified Au or Pt electrodes was performed by CV

a t 5 mV s-1 using 5 mM ferricyanide solution in 0.1 M

potassium chloride.

On-chip reference electrode preparations were undertaken by

modification of the gold pseudoreference electrodes by elec-

trodeposition of silver from an aqueous 5 mM silver nitrate

solution in 50 mM potassium nitrate and 0.5 M potassiumthiocyanate. Electrodeposition of silver was achieved at a fixed

potential of -0.15 V (vs a Ag wire) for 10 min. Silver/silver

chloride (Ag/AgCl) was formed by the immersion of the on-chip

silver electrode in 1 M iron(III) chloride for 60 s.

Surface Modification. Surface modification of the chips was

performed as follows: (i) pretreatment of the chips with oxygen

plasma (150 W, 10 min); (ii) silanization by immersion of the chips

in 3% 3-aminopropyltrimethoxysilane (APTES) (Gelest) in a 19:1

dilution of methanol/deionized (DI) water, followed by washing

with methanol and DI water; (iii) heat curing of the chips at 120

C for 15 min; (iv) deposition of a cross-linker by immersion of

the chips in dimethylformamide (DMF) containing 10% pyridineand 1 mM 1,4-phenylene diisothicyanate (PDITC) (Fluka) for 2 h;

(v) final washing of the chips with DMF and 1,2 dichloroethane

followed by drying under a stream of nitrogen. The primary

immunoreagents of the sensor were covalently immobilized

following this stage of the surface modification (Antibody Im-

mobilization onto the Chip Device).

Characterization and assessment of the surface functionaliza-

tion was undertaken using fluorescently tagged ssDNA (single

stranded DNA with an amine anchor on the 5 end, for attachment

to the PDITC cross-linker, and a fluorescent tag on the 3end)

(All from Sigma-Proligo). A 20 M solution of the DNA was diluted

1:5 in printing buffer (1 M Tris-HCl (pH 7) with 1% v/v N,N-diisopropylethylamine) and deposited onto the chip surface (either

across the whole surface or as a spot deposition over the working

electrode area). The chips were then incubated overnight at 37

C in a dark, humid chamber. Unreacted cross-linker moieties

were capped by immersion of the chips in 50 mM 6-amino-1-

hexanol and 150 mM N,N-diisopropylethylamine in DMF for 2 h,

followed by washing with DMF, MeOH, and DI water. Surface

coverage of the modified chip surface with the fluorescent DNA

was assessed using a Ziess Axiscope fluorescent microscope.

Surface modification of both microsquare electrode arrays and

microband electrode arrays was undertaken for comparative study,

as they both have identical surface chemical properties and the

only difference was in their electrochemical signals (due to

differing diffusion profiles, which are dependent on the electrode

shape and size). This parallel work placed less restriction on

available surfaces for modification studies.

Antibody Immobilization onto the Chip Device. The

capture antibody (anti-aflatoxin M1) was diluted (96 g mL-1)

with carbonate buffer (0.1 M, pH 9.6), of which 1 L of the

antibody solution was placed onto the device. These were

stored overnight at 4 C in humid conditions to allow covalent

attachment via the PDITC cross-linker. The devices werewashed twice with 10 mM PBS-T pH 7.4 buffer, once with water

using a dispensing bottle, and then shaken dry. After the

devices were dried, 3 L of 0.1% NH4OH in water was added

for 60 min at room temperature to deactivate any unreacted

PDITC cross-linker and then washed and dried. A volume of 1

L of 40 g mL-1 anti-aflatoxin M1 antibody was placed onto

the devices and incubated at 37 C for 2 h in humid conditions.

The electrode arrays were then washed and dried as reported

above and stored at 4 C until used.

Assay Development for the Chip Device. For the optimiza-

tion of TMB electrochemical detection using the MEA, differential

pulse voltammetry was employed. The working MEA with theimmobilized PDITC cross-linker was first capped using 1% NH4OH

at room temperature for 1 h before 0.5 mM TMB in 10 mM

citrate buffer and 0.1 M KCl was placed onto the electrode

surface. The electrode array was connected to an AUTOLAB

potentiostat (Eco chemie, The Netherlands) via a custom-made

connector and the potential was scanned from -0.5 to 0.5 V

using the on-chip pseudoreference and counter electrodes.

The immunoreaction of aflatoxin M1 to the activated electrode

surface was achieved by placing 1 L of sample or standard,

mixed 1:1 with aflatoxin M1-HRP (diluted 1:10 with 10 mM

PBS, pH 7.4) onto the antibody immobilized MEA and

incubated at 37 C for 120 min. The devices were washed twicewith 10 mM PBS-Tween (0.05% Tween 20) pH 7.4 buffer and

once with water using a dispensing bottle and then was shaken

dry. The bound HRP-conjugate was then determined using a

TMB/H2O2 solution. This solution was prepared by dissolving

1 mg of TMB in 150 L of DI water, and 20 L of this stock

solution was mixed with 2 L of 30% hydrogen peroxide and

made up to 1 mL using 10 mM citrate buffer (pH 5.2)

containing 0.1 M KCl at 37 C. A 4 L aliquot of the TMB/

H2O2 solution was placed onto the MEA immediately prior to

analysis. The stock solution of TMB was prepared daily and

stored in the dark prior to use.

The electrochemical measurements were performed by con-necting the microarray to the AUTOLAB potentiostat. A condition-

ing prepotential was applied first for 5 s at a potential of +268

mV and then the potential was set to +168 mV for measurement

(5 min). Preconditioning the electrode as reported above before

data collection has been shown previously to increase the signal

achieved from the immunoassay.34,42,43 Samples of full fat milk

were pretreated by centrifugation at 9000 rpm (5 min), and an

aliquot was taken from below the upper fat layer and used in the

analysis. Curve fitting of data reported in this paper was carried

out using Graphpad Prism version 5.02 from Graphpad software.

Surface Analysis of the Microelectrode Array. The surface

of the MEAs was characterized by atomic force microscopy (AFM)5293Analytical Chemistry, Vol. 81, No. 13, July 1, 2009

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and scanning electron microscopy (SEM) to monitor the im-

mobilization of the antibody. The SEM images were taken using

a Philips XL30 scanning field emission gun (SFEG, U.K.). The

AFM images were obtained in a wet environment using a

Dimension 3000, from Digital Instruments (now Veeco Instru-

ments, U.K.). The AFM tips used were silicon probes used in the

tapping mode. The probes were 225 m 38 m 7 m with a

typical resonant frequency of 160 kHz. The scan speed applied

was between 0.5 and 1 Hz.

The surfaces of two MEA devices were analyzed in detail using

AFM to monitor immobilization of the antibody. One of these

sensors surface was prepared by immobilizing the capture

antibody prior to the surface analysis. A volume of 1 L of 96 g

mL-1 of capture antibody (Pierce, U.K.) was placed onto the

MEA surface at pH 9.6 and incubated at 4 C overnight. The

surface was washed with 10 mM PBS-T and H2O, and then

the excess linker compound was deactivated using 4 L of 0.1%

NH4OH for 1 h at 37 C.

Safety Awareness. All laboratory glassware and consumables

which had been contaminated with aflatoxin M1 were stored

overnight in 5% sodium hypochlorite, and then acetone was

added to make the solution 5% acetone by volume. The

decontamination solution was then stored at room temperature

for a minimum of 30 min before disposal.

RESULTS AND DISCUSSION

Cell-on-Chip Device Characterization. Each gold cell-on-a-chip device consisted of gold counter and reference electrodes

with a gold MEA working electrode (Figure 1). Each array

consisted of 35 microsquare electrodes with 20 m 20 m

dimensions and edge-to-edge spacing of 200 m (an electrode

width to spacing ratio of 10). These dimensions and spacings were

chosen to avoid overlapping diffusion layers between neighboring

electrodes in the array.

The surface of the sensor was first characterized using SEM

imaging. These were taken using sFEG rather than a conventional

SEM since the sFEG gives better resolution. With the use of a

high-resolution scanning electron microscope (sFEG) at a low

magnification, images (

80 magnification) of the working MEA

for the microsensors were taken. Figure 1b, clearly shows the

layout of the working MEA with the 35 elements etched into the

surface and also the image of the single recessed electrode,

obtained using AFM. This design was used for the development

of the immunosensor.

Electrochemical characterization of the cell-on-a-chip micro-

electrodes was conducted by CV using 1 mM FcCOOH in PBS at

5 m V s-1

. Characterization of the working electrode withexternal counter and reference electrodes was undertaken to

ensure that the electrodes were functioning correctly. Further

characterization of the cell-on-a-chip devices was undertaken

following modification of the on-chip Au reference electrode

with Ag and with Ag/AgCl (Figure 2). An apparent shift in the

redox potential of the electroactive species can be seen on

changing the reference electrode from the pseudoreference

unmodified gold to the Ag and Ag/AgCl reference electrodes. The

latter modification results in a similar redox potential to the

external Ag|AgCl reference electrode. All CVs show characteristic

steady-state cyclic voltammograms as expected for microelec-

trodes with sufficient interelectrode spacing to achieve indepen-

Figure 1. (a) The three-electrode chips were fabricated with one working electrode area (35 electrodes in the array), a counter electrode, and

a reference electrode area. (b) The whole working microelectrode of the untreated surface at 80 magnification using a sFEG. Atomic forcemicroscopy image of a single element of the array for the untreated working microelectrode (image 40 m 40 m).

Figure 2. Characterization of Au cell-on-a-chip microelectrodes with1 mM FcCOOH in 0.01 M PBS using cyclic voltammetry at 5 mV s-1.

The influence of different on-chip reference electrode preparations

with working electrodes consisting of 35 Au microsquares with 20

m 20 m dimensions and edge-to-edge spacing of 200 m.

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dent diffusion profiles for each element of the array and gave

currents in agreement with established models for diffusion-

controlled currents at microelectrodes.8-10 Since the choice of

the reference or pseudoreference electrode did not influence the

magnitude of the current, only its position on the potential axis,

the use of a gold pseudoreference electrode for subsequentmeasurements was chosen. This simplified the electrochemical

device preparation by avoiding the necessity to prepare the on-

chip Ag and Ag/AgCl electrodes.

Device Surface Modification. Where small arrays of micro-

electrodes are used, the total electrode surface area is relatively

small, and thus for biosensor applications, a limited amount of

biorecognition material can be immobilized. This further limits

the information generating capacity of the sensor and compro-

mises the sensitivity. In order to improve the biocapture capacity

and sensitivity of the chips, modification of the sensor insulating

area (silicon nitride) around the microelectrodes was investigated.

Aside from the larger immobilization area, a further advantage ofmodification to this area is that the electrode surface remains less

affected by the surface modifications, thus the signal is less

attenuated by the immobilized reagents.

The modification of the chips using the amino-silane anchor

(APTES) and cross-linker (PDITC) was carried out on the sensors

surfaces. The chemistries were applied to the whole chip surface

and the surface coverage assessed using fluorescently tagged

DNA. These studies were also conducted with gold microband

electrodes consisting of four 50 m 500 m bands with 500 m

edge-to-edge separation and platinum microsquare electrodes for

comparison of sensor performance. The modification provided a

more hydrophobic surface, hence spot depositions could be moreprecisely and reproducibly applied to the chip. Figure 3 shows

the fluorescent microscopic images, demonstrating successful

immobilization of the DNA onto the silicon nitride surface.

Immobilization was investigated by coverage of the entire chip

surface, i.e., the whole of the three-electrode cell-on-chip surface

(in this case examined using Au microband electrodes, Figure

3a), and by spot deposition of the DNA over the working electrode

area only (tested using the Pt microsquare electrodes, Figure 3b).

The edge of the spot can be clearly seen in Figure 3b, indicating

successful modification of the surface in that area of the chip.

However, it is less clear from these fluorescence images if the

metal electrode surfaces were also modified.

The aim of the surface modification was for attachment of the

biorecognition material mainly on the surrounding silicon nitride

layer; therefore, an assessment of the surface coverage on the

gold and platinum electrode surfaces was carried out by voltam-

metry. The most widely used method to assess the extent of

electrode surface coverage by DNA is based on cyclic voltammetryof ferricyanide.31 Shielding of the ferro-/ferricyanide ions by the

immobilized DNA can be attributed to a combination of physical

coverage of the electrode by the DNA and electrostatic repulsion

between the negatively charged redox couple ions and the DNA

phosphate backbone.32 Figure 4 shows the cyclic voltammetry of

ferricyanide using the DNA-modified gold and platinum MEAs.

For each chip, the sensor response was assessed using the on-

chip reference and counter electrodes and also using external

reference and counter electrodes. The response of the DNA-

modified platinum electrodes shows a reduced current with

respect to the unmodified platinum, suggesting some modification

of the electrode surface as well as the surrounding silicon nitridelayer around them (Figure 4b). The modified gold electrodes, in

contrast, show that the current has not been decreased with

respect to the unmodified electrodes, suggesting that the modi-

fication procedure has not significantly covered the gold surface

(Figure 4a). This is attributed to the reaction of the silane reagent

APTES with residual oxides on the surface of the platinum leading

to DNA immobilization on the platinum. Gold cell-on-chip sensors

were used for subsequent immunosensor development including

the use of the on-chip gold pseudoreference electrode.

Immunoassay Development. For the detection of aflatoxin

M1, a monoclonal antirat antibody was employed as the sensing

molecule for the microarray electrode. In order to ensureoptimal orientation of the monoclonal antibody on the sensor

surface, a polyclonal capture antibody was covalently im-

mobilized to the microarray through the PDITC cross-linker

chemistry, since covalent immobilizations can cause the

antibody to lay in a side on orientation33 and therefore results

in poor binding efficiency to the analyte. By incorporation of a

capture antibody, the sensing antibody becomes highly

efficient. With the monoclonal sensing antibody immobilized

to the surface of the microarray, the detection of aflatoxin M1

was carried out by a competitive reaction between the free

aflatoxin M1 in the sample and an aflatoxin M1-horseradish

peroxidase conjugate. This assay procedure was developed and

Figure 3. Fluorescence microcope images of surface modified chips using silanization/PDITC chemistry with fluorescently tagged DNA: (a)

gold microband electrodes and (b) platinum microsquare electrodes showing the microsquare electrode opening in the silicon nitride layer,underlying Pt electrode, and edge of spot deposition.

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characterized in our previous work using screen-printed sen-

sors (macrosensors) for aflatoxin M1 analysis34 before it was

transferred after minimal modification to the surface of themicroelectrode sensor.

The developed method for the macrosensors34 utilized passive

adsorption onto a carbon electrode surface as an immobilization

protocol. The immobilization method for the microsensors was

by covalent attachment using the amino-silane anchor and PDITC

cross-linker, which was a deviation from the macrosensor method.

The validation of successful immobilization of the antibody onto

the surface was performed using atomic force microscopy. After

the immobilization of the assay reagents on the microelectrode,

a 3D image of the surfaces inside a microsquare electrode was

then conducted. This was compared to bare array surface as

shown in Figure 1. The results from Figure 5 indicate that the

immobilization of the antibodies cause a change of the root mean

square (rms) roughness from 1.27 to 2.37 nm (increase of 0.84

nm) agreeing with the observation of other researchers35-37 that

upon the addition of protein to a sensor surface the roughness

increases. Both the surface roughness and topography indicated

quantifiable differences, and visual evidence was seen using 3D

imaging and phase control. This shows that the antibodies have

also passively adsorbed to the gold working MEA while being

covalently immobilized to the silicon nitride surrounding. It is

estimated that these surface-adsorbed antibodies will have a

minimal enhanced effect on the final results due to the large

surface of the silicon nitride and small surface of the gold MEA.

To eliminate this passive adsorption, the gold microarray surface

may need to be blocked using short chain alkane thiols. However,

this is not necessary in this work as the effect is small.

With the use of the immobilization protocol, several devices

were prepared for aflatoxin M1 measurements in buffer as wellas spiked commercial milk. Before electrochemical quantifica-

tion of aflatoxin M1 could be demonstrated, the electrochemical

test parameters were optimized for the electrochemical MEA

devices. The electrochemical detection of TMB is achieved

using either the reduction peak or oxidation peak at+100 mV

vs Ag/AgCl38-40 and -100 mV vs Ag/AgCl,41 respectively.

Differential pulsed voltammetry was employed to find the maxi-

mum detection potential of TMB using the gold microelectrode

array. The voltammogram (data not shown) showed that the

maximum peak signal occurred at a potential of+168 mV (vs the

gold pseudoreference electrode); therefore, the detection of TMB

using chronoamperometry was set at+

168 mV (vs the goldpseudoreference electrode) and a preconditioning potential of 100

mV above the detection potential (+268 mV vs the gold pseu-

doreference electrode) was applied for 5 s before chronoampero-

metric measurement.42,43The coefficient of variation (%CV) value

for current obtained using modified chips (with the amino-silane

anchor (APTES) and cross-linker (PDITC)) from TMB detection

using chronoamperometry analysis with three replicate electrodes

) 4%. With the use of these settings, a calibration graph for

aflatoxin M1 was determined in buffer as well as in spiked milk

samples using the microarray immunosensor as shown in

Figure 6. Error bars indicate the standard deviation (n ) 3).

The results (Figure 6a) show that the microelectrodes are very

sensitive and are able to detect levels of toxins lower than the

current legislative requirements of 50 ng L-1.25 The analytical

sensitivity in buffer is less than 1 ng L-1 (calculated as the

amount of aflatoxin M1 needed to produce a 25% decrease in

the signal).47 The r2 value for the curve was 0.986 using the

Graphpad Prism one site - fit Ki curve fitting function. With

the devices performing well in pure buffer solutions, further

examination was carried out to assess the performance in a

milk matrix. With the use of the pretreatment for the milk

samples described in the Experimental Section, a calibration

curve was established using spiked milk samples. Figure 6b shows

that the dynamic detection range in milk samples occurred

between 10 and 100 ng L-1 and the analytical sensitivity(calculated as the amount of aflatoxin M1 needed to produce a

(35) Parra, A.; Casero, E.; Pariente, F.; Vazquez, L.; Lorenzo, E. Sens. Actuators,

B 2007, 124, 3037.(36) Tsai, Y. C.; Huang, J. D.; Chui, C. C. Biosens. Bioelectron. 2007, 22, 3051

3056.(37) Vianello, F.; Zennaro, L.; Rigo, A. Biosens. Bioelectron. 2007, 22, 2694

2699.

(38) Badea, M.; Micheli, L.; Messia, M. C.; Candigliota, T.; Marconi, E.; Mottram,

T.; Velasco-Garcia, M.; Moscone, D.; Palleschi, G. Anal. Chim. Acta, 2004,

520, 141148.(39) Fanjul-Bolado, P.; Gonzalez-Garca, M. B.; Costa-Garca, A. Anal. Bioanal.

Chem. 2005, 382, 297302.(40) Butler, D.; Pravda, M.; Guilbault, G. G. Anal. Chim. Acta 2006, 556, 333

339.(41) Micheli, L.; Grecco, R.; Badea, M.; Moscone, D.; Palleschi, G. Biosens.

Bioelectron. 2005, 21, 588596.(42) Lu, H.; Conneely, G.; Pravda, M.; Guilbault, G. G. Steroids 2006, 71, 760

767.(43) Conneely, G.; Aherne, M.; Lu, H.; Guilbault, G. G. Anal. Chim. Acta 2007,

583, 153160.

Figure 4. Electrochemical characterization of DNA-modified and

unmodified chips with 5 mM FeCN in 0.1 M KCl: (a) gold microbandelectrodes and (b) Pt microsquare electrodes. Data are based on

cyclic voltammetry of three separately modified chips (a-c) of Auand Pt with comparison between on-chip and external reference

electrodes. Scan rate 5 mV s-1.

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25% decrease in the signal)47 at 8 ng L-1 which surpasses

previous reported immunosensors for aflatoxins.

38,41,44-48

The

r2 value for the curve was 0.999 using the graphpad prism one

site - fit Ki curve fitting function.The different response curve shown for analysis in buffer and

then in milk is due to the use of different batches of sensors to

conduct the analysis which at the time of the experiments were

manually assembled. Improvements in assembly will minimize this

variability, although clearly on-chip variation is acceptable, as

shown in the response curves produced in milk samples. The %CV

for the immunoassay on the MEA (n ) 3) was 6%.

In our previous work for the development of macrosensors,34

milk was found to interfere with the electrochemical measurement

using the carbon screen-printed electrode. It was deduced that

whey proteins, with a specific focus on R-lactalbumin, were

blocking the carbon electrode surface. The macrosensors were

therefore blocked using a PVA layer; additionally milk samples

required pretreatment with calcium chloride to reduce whey

proteins blocking of the electrode surface. With the current

protocol utilizing the MEAs, the sensor chips were not blocked

by any blocking agent as with the carbon based screen-printed

electrodes, no matrix interference was observed, and no pretreat-

ment with calcium chloride was required. For the gold MEAs,

blocking of the surface with polymers is not required due to the

lower absorptive properties of the gold compared to carbon and

perhaps the lower surface area (due to smoothness) of litho-

graphically produced electrodes over screen-printed technology

which produced a comparably rough surface.

As the research reported here is an initial investigation intodeveloping new technology in food safety testing, the work is still

in progress and gives proof of principle of the technology for a

new and important application. The idea is for testing performed

(44) Carlson, M. A.; Bargeron, C. B.; Benson, R. C.; Fraser, A. B.; Philips, T. E.;

Velky, J. T.; Groopman, J. D.; Strickland, P. T.; Ko, H. W. Biosens.

Bioelectron. 2000, 14, 841848.(45) Nasir, M. S.; Jolley, M. E. J. Agric. Food Chem. 2002, 50, 31163121.(46) Gaag, B.; Spath, S.; Dietrich, H.; Stigter, E.; Boonzaaijer, G.; Osenbruggen,

T.; Koopal, K. Food Control 2003, 14, 251254.(47) Ammida, N. H. S.; Micheli, L.; Palleschi, G. Anal. Chim. Acta 2004, 520,

159164.(48) Chiavaro, E.; Caccgiolo, C.; Berni, E.; Spotto, E. Food Addit. Contam. 2005,

22, 11541161.

Figure 5. (a) The surface roughness taken inside of the untreated element, analyzed by atomic force microscopy. (b) The surface roughness

inside a treated element, analyzed by atomic force microscopy. For the treated microelectrode, 1 L of 96 g mL-1 dilution of capture antibodywas immobilized at 4 C overnight, and then excess linker compound was deactivated using 4 L of 0.1% NH4OH for 1 h at 37 C and AFM

images were taken immediately for the gold microsquare array.

Figure 6. Standard curve for aflatoxin M1 detection using thedeveloped MEA cell-on-chip device (a) in buffer solution and (b) in

milk. A volume of 1 L of the sample or standard + aflatoxin M1 HRP(diluted 1:10 with PBS) was placed and incubated at 37 C for 120

min before detection with TMB (0.5 mM) and hydrogen peroxide (1

mM) in citrate phosphate buffer pH 5.2 with 0.1 M KCl at +168 mV.Error bars indicate the standard deviation (n) 3).

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at the farm and as such only full fat commercial fresh milk was

tested in this investigation before farm testing can be performed.

Future work will investigate analysis of a range of milk types and

products. Further sensor optimization is required before a final

device can be considered for commercialization, such as the

reduction of incubation time of the assay, and stabilization of

the immunoreagents on the sensor surface, such as drying the

immuno-components with an osmolyte e.g., trehalose or betaine.

Few comparable publications for MEA-based immunosensors

are available. The most directly comparable report in terms ofmethod and application is that reported by Dill49 who has

produced an immunosensor for R1 acid glycoprotein with a

detection limit of 5 ng L-1. The device reported here has a

detection capability of 8 ng L-1 and operates in competitive

immunoassay format. With review of other electrochemical

immunosensors for the detection of aflatoxin M1, this surpasses

the work of Micheli et al., who achieved limits of detection at

25 ng L-1 using screen-printed technology.41

CONCLUSIONS

Gold MEAs (35 elements) with on-chip reference and counter

electrodes, fabricated using photolithographic techniques, wereused in this investigation to develop an immunosensor for aflatoxin

M1 in milk samples. The microarray was characterized using

cyclic voltammetry and showed good performance similar to

that reported in the literature for MEA devices. Surface

modification was successfully conducted for the covalent

immobilization of the capture antibody. A competitive immu-

noassay was then performed on the surface of the sensor using

the TMB/H2O2 electrochemical detection system with HRP as

the enzyme label. The sensors were reproducible and very

sensitive giving a detection limit of 8 ng L-1 in milk matrix.

The sensor has been designed for one use (disposable);

however, our current unpublished work indicates that the chip

can be recycled after plasma cleaning. Further optimisation of

the assay protocol can result in a lower detection limit. Theimprovements seen by employing gold microelectrodes rather

than screen-printed electrodes for aflatoxin M1 detection

indicate that for high sensitivity applications, such as low

detection limits, a boundary has been surpassed allowing for

development of many new applications which have previously

only been achieved using elaborate instrumentation.

ACKNOWLEDGMENT

The authors thank the European Commission for supporting

this work (Project FP6- IST1-508774-IP GOODFOOD: Food safety

and quality with microsystems technology). Damien W. M.

Arrigan thanks the Science Foundation Ireland for ongoingsupport (Grants 02/IN.1/B84 and 07/IN.1/B967).

Received for review March 10, 2009. Accepted May 11,2009.

AC900511E(49) Dill, K.; Montgomery, D. D.; Ghindilis, A. L.; Schwarzkopf, K. R.; Ragsdale,

S. R.; Oleinikov, A. V. Biosens. Bioelectron. 2004, 20, 736742.


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