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Special Lecture SeriesBiosensors and Instrumentation

Lecture 3: Introduction to Electrochemical BiosensingElectrochemical biosensors are integrated receptor-transducer devices – capable of providing selective and quantitative analytical information using a biological recognition element. A biomolecule is immobilized onto an electrode that has been modified by coating it with an electronically conducting, semiconducting, or ionically conducting chemical. Coating materials that have been used include electrochemically prepared polyaniline and polypyrrole films, or other conducting polymers. A biosensor can be used to monitor either biological or nonbiological chemical analytes. They can also be used for the detection of microorganisms. Suitable receptor-transducer systems for this are bacteria, yeast, or oligonucleotide probes coupled to electrochemical, piezoelectric, optical or calorimetric elements.Electrochemical enzyme-based biosensors are widely used because of their practical advantages, which include their low fabrication cost, simplicity of operation, high selectivity, and their ability to perform real-time detection. There are three main types of Electrochemical Biosensors:

• Potentiometric biosensors use ion selective electrodes to determine changes in the concentration of chosen ions.

• Amperometric biosensors measure the electric current associated with electron flow resulting from redox reactions.

• Conductometric biosensors measure changes in the conductivity of a medium as a result of enzyme reactions that change its ionic composition.

They are readily miniaturised using the microfabrication techniques developed for the electronics industry. This includes their incorporation into field-effect devices (e.g., ISFETs), screen-printing, and ink-jet printing for depositing nano-litre volumes of reactants and for laying down tracks of conducting ink.

Potentiometric SensorsThe basic principle behind potentiometric sensor measurements is the development of a voltage related to the analyte activity (concentration) [A] in the sample through the Nernst relation:

Potentiometric sensors will generally require a reference electrode as well as the indicator (working) electrode to be in contact with the test sample solution. The use of ion-selective membranes can make these sensors sensitive to various ions (e.g, hydrogen, fluorine, iodine, chlorine ions) in addition to gases such as carbon dioxide and ammonia. Enzyme systems, that change the concentration of any of these ions or gases, can also be incorporated into the sensor in order to be able to measure enzyme substrate concentrations, or to detect inhibitors (e.g., heavy metal ions, insecticides) or modulators of the enzyme.

E = Eo +RT

nFln [A]

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Ideally, the potential difference between the indicator and reference electrode is proportional to the logarithm of the ion activity or gas fugacity. However, this is only the case when:

• The membrane or indicator electrode surface layer is 100% selective for the test analyte, or

• There is a constant or low enough concentration of interfering ions, and

• Potential differences at various phase boundaries (junction potentials) are either negligible or constant, except at the membrane-sample solution interface.

Many of these sensors take the form of a pH electrode to measure the activities of enzymes (and hence the concentration of the specific substrate for that enzyme) which produce or consume protons as a result of catalysis. Examples of enzymes that can be used in this way are urease, glucose oxidase, penicillinase and acetylcholinesterase - to monitor the concentrations of urea, glucose, penicillin, and the neurotransmitter acetylcholine (or some pesticides that inhibit acetylcholinesterase) respectively.

Strengths:

• A wide concentration range for detection of ions (typically 1 µM to 0.1 M).

• Can perform continuous measurements (ideal for clinical/environmental use).

• Inexpensive and portable.

Weaknesses:

• pH buffers are often required to maintain optimum enzyme activity, and this can limit the dynamic range of detection of the analyte for enzyme-based sensors.

Amperometric: These are the most commonly reported class of biosensor. They typically rely on an enzyme system that catalytically converts electrochemically non-active analytes into products that can be oxidized or reduced at a working electrode. This electrode is maintained at a specific potential with respect to a reference electrode. The current produced is linearly proportional to the concentration of the electroactive product, which in turn is proportional to the nonelectroactive enzyme substrate. Enzymes typically used in amperometric biosensors are oxidases that catalyze the following class of reactions:

Substrate + O2 → Product + H2O2 1

As a result of the enzyme-catalyzed reaction, the substrate (analyte) concentration can be determined by amperometric detection of oxygen or hydrogen peroxide (H2O2). An example of this configuration would be an oxygen-consuming enzyme coupled to an oxygen-sensing electrode. The ambient oxygen concentration is then continuously monitored as it diffuses through a semi-permeable membrane and is reduced at a platinum (Pt) electrode. Other common configurations include the use of oxidases specific to various substrates to produce H2O2.

Please note we’re using a definition of “substrate” that might be unfamiliar to you as electronics engineers. 1

The “substrate” here is any substance on which an enzyme acts. So for the glucose sensor, the substrate of the enzyme Glucose Oxidase is glucose

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During measurement, the working electrode may act as an anode or a cathode, according to the nature of the analyte. For example, a glucose-sensitive biosensor that uses glucose oxidase could detect the H2O2 produced by the enzymatic reaction by polarising the working electrode to a positive potential (+0.6V vs. SCE), or by polarising the working electrode to a negative potential (-0.65V vs. SCE) to monitor oxygen.

These sensors which use an oxidase enzyme and detect the products electrochemically are considered to be first generation devices which have a number of problems.

Strengths:

• The use of low-cost and disposable electrodes.

• High degree of reproducibility that is possible for these (one-time use) electrodes eliminates the cumbersome requirement for repeated calibration.

• The instrumentation for these biosensors is inexpensive and compact, allowing for the possibility of on-site measurements.

Weaknesses:

• Tend to have a small dynamic range due to saturation kinetics of the enzyme.

• Potential interference to the response if several electroactive compounds can generate false current values. (These effects have been eliminated, for clinical applications, through the use of selective membranes, which carefully control the molecular weight of the charge of compounds that have access to the electrode.)

• Require oxygen for the enzyme activity to transfer electrons. If the oxygen content of the measured solution is too low then the reaction rate will depend on this rather than the glucose concentration.

One method to get round these issues involves the use of membranes as mentioned above. The outer membrane can control the flux of oxygen and glucose between the sensed environment and the immobilised enzyme layer. This can prevent the sensor being dependent on O2 concentration or on the diffusion barrier in an unstirred solution. The inner barrier is designed to prevent interferents which are electroactive at the same potential as H2O2 from reaching the electrode. Most of these interferents are larger molecules which enables this discrimination.

Other solutions to cross tolerance with interferants in H2O2 sensing may involve the use of other materials, enzymes etc. to lower the potentials used and avoid other electroactive substances. The paper mentioned on glucose sensing earlier in the course uses “Prussian Blue” immobilised in chitosan which has a dual effect of catalysing the peroxide sensing and blocking other molecules [1].

Amperometric biosensors operate by measuring the current generated by oxidation or reduction of redox species at an electrode surface, which is maintained at an appropriate electrical potential. The current observed has (hopefully) a linear relationship with the concentration of the analyte. However, the direct electron transfer between the redox-active site of an enzyme immobilized at an electrode surface is normally prohibited by an intervening, insulating, part of its polypeptide structure. Some kind of charge carrier has then to be used as an intermediate between the enzyme redox centre and the electrode. In glucose

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oxidase the redox-active site is a flavin adenine dinucleotide (FAD) molecule at the heart of the protein (slide 18).

Mediated Amperometric Biosensors: If a redox enzyme resists direct electrical communication with an electrode, this can be overcome using synthetic or biologically active charge carriers as intermediates between the redox centre and the electrode. An example of this is shown schematically in slide 20 using potassium ferrocyanide [K4Fe(CN)6] whose anion [Fe(CN)6]4– (known as ferrocyanide) readily oxidizes to ferricyanide:

[Fe(CN)6]4– ↔ [Fe(CN)6]3– + e–

The ferri/ferrocyanide redox couple has been used in glucose sensors as an intermediate between the redox centre of glucose oxidase and the working electrode, for example. Other mediators commonly used are phenylene-diamine, diaminobenzidine, hydroquinone and aminosalicylic acid. Such electrochemical mediators must possess a number of properties:

• They must react rapidly with the reduced form of the enzyme,

• They must be sufficiently soluble, in both their oxidised and reduced forms, to be able to diffuse rapidly between the active site of the enzyme and the electrode surface,

• The reduced form of the mediator should not readily react with oxygen,

• They should be stable in both their reduced and oxidised forms and should not react with other chemicals in the sensor,

• It must be able to compete with the enzyme’s natural substrate (e.g., molecular oxygen in the case of oxidases),

• Its redox potential should provide an appropriate potential gradient for electron transfer between an enzyme active site and an electrode,

• It should exhibit reversible electrochemistry, having a large rate constant for the interfacial electron transfer at the electrode surface.

Unmediated Amperometric Biosensors: These are biosensors where direct electron transfer between the redox-active site of an immobilised enzyme at an electrode surface is possible. An example of direct electron transfer between an enzyme and an electrode is depicted in Slide 23. Slide 24 shows schematically the direct electrical coupling of a peroxidase enzyme (e.g., horse radish peroxidase) with an electrode, for the detection of hydrogen peroxide (H2O2) produced by an oxidase enzyme that can make a direct, nonmediated, redox contact with the peroxidase. The peroxidase catalyses the following reaction:

H2O2 + 2e– + 2H+ → 2H2O

Direct electron communication between enzyme redox sites and electrodes can be facilitated by changing the surface morphology of the electrodes at the nanoscale – as for example by coating the surface with metal nanoparticles.

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One of the reasons for exploring un-mediated and next generation glucose sensors is for applications in implantable sensor technology where mediators may not be acceptable as they are often considered to be toxic or unstable in the environment of the human body. Continuous glucose sensors that can measure for many days/weeks/months are extremely desirable and the first attempts at delivering these are starting to come onto the market. We will look at the development of implantable sensors and other microsystems for biomedical applications later in the course.

Conductometric: These sensors measure the conductivity of a medium at a series of electrical frequencies. AC measurements allow conductivity changes to be determined whilst minimising undesirable electrochemical processes. Conductance is directly related to the amount of ions in a medium, and since many enzyme-linked reactions result in a change in ion concentration they are suitable for conductometric biosensors.

Enzyme reactions that produce or consume ionic species depend on the total ionic strength of the medium and changes in its conductance/capacitance can be relatively small. Various planar interdigitated electrode configurations have been devised as the conductometric transducer. The interdigitated geometry provides long electrode tracks (~ 1 cm) within a small sensing area (0.2 sq. mm). The electrodes can be fabricated by depositing platinum paste or silver-palladium electrodes onto an insulating substrate, or by the vacuum evaporation of metal electrodes and photolithography.

Ion-sensitive organic conducting films can be deposited onto the electrode array as the sensing element. An example is the use of phlalocyanine in peroxidase-linked immunoassays, in which the peroxidase converts iodine ions to iodine molecules. Sensing enzymes can be immobilised onto the electrodes in a paste or gel form. An example of this is a urea sensor, using the enzyme urease to catalyze the hydrolysis of urea and produce ionic species (ammonium, bicarbonate and hydroxyl ions):

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Strengths:

• Inexpensive, reproducible and disposable.

• The interdigitated electrodes can be fabricated using photolithography - hence miniaturised sensors can be mass produced and used in multi-sensing and portable devices.

Weaknesses:

• The ionic species produced must significantly change the total ionic strength to obtain reliable measurement.

• This requirement can increase the detection limit to unacceptable levels and results in potential interferences from variability in the ionic strength of the sample.

A high level of detection sensitivity can be achieved using modified bridge circuits and high input resistance operational amplifiers whose outputs are directly proportional to the sensor conductance.

urea + 3H2Ourease! 2NH+

4 +HCO3 +OH

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Impedimetric Biosensors Similar interdigitated electrodes may also be applied to capacitive measurement of biosensor systems where the dielectric permittivity between the electrodes is altered by a bio-recognition event. An example might be an immunosensor where antigens are immobilised in the spaces between the electrode fingers. Specific binding of antibodies will alter the overall permittivity leading to a change in the capacitance that can be measured. As with the conductometric sensors described above, the electrodes are measured with a low amplitude AC signal at a DC potential which is chosen in order to avoid any Faradaic electrochemical reaction at the electrode surfaces. In a purely capacitive sensing system it may even be possible to apply a dielectric layer over the electrodes in order to passivate them and ensure that no chemical reaction can take place between the electrodes and the solution being measured. This may also be required in order to present a desirable surface for binding of the biosensing molecules. However, this passivation layer would have to be chosen so that it did not significantly reduce the sensitivity of the measurement. In [2], Katz and Willner describe a variety of biosensors with conductometric or impedimetric operation. This article also looks at impedance based biosensors which make use of Faradaic electrochemical processes. This is the subject of the next part of the course which is focussed on electrochemical impedance spectroscopy.

Electrochemical Impedance Spectroscopy Impedance spectroscopy can be used to analyze electrical processes occurring in a system. The technique is particularly sensitive to changes in both surface and bulk effects, and as such is a valuable technique to apply for electrochemical research and sensor applications.A good example to consider is a simple reversible reaction at a working electrode of the form:

where both the oxidised and reduced species, Ox and R, are soluble in the aqueous solution at the electrode. An impedance spectrogram can be obtained by applying an AC voltage to the electrochemical cell. The resulting current-potential response is given by the Butler-Volmer equation (equation 8 of Lecture 2):

(1)Expansion of the exponential functions in this equation into their series forms gives:

!

!

For small perturbations of the working electrode’s potential ([E – Eo] << [RT/αnF]) the quadratic and higher terms can be ignored, and for αC = αA = 0.5, we have:

(2)

Ox + ne $ Red

I = Io

exp

↵A

nF (E Eo

)

RT

exp

↵

C

nF (E Eo

)

RT

I = Io

"1 +

↵A

nF (E Eo)

RT+

1

2!

↵A

nF (E Eo)

RT

2

+ . . .

#

Io

"1 ↵

C

nF (E Eo)

RT+

1

2!

↵C

nF (E Eo)

RT

2

. . .

#

I = Io

nF

RT(E Eo)

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with

, or

!

depending on whether the potential is perturbed to a value above or below Eo, respectively. Thus, if the potential of the working electrode is perturbed by just a few millivolts (≤ 5mV) about the value of the equilibrium potential Eo, we can assume that the current-potential response is approximately linear. The ratio (E – Eo)/I has dimensions of resistance, and the concept of a charge-transfer resistance Rct can be introduced and defined as:

(3)This resistance component is small when the exchange current Io is large (corresponding to a high charge transfer rate). As discussed in Lecture 2, apart from the impedance to current flow related to charge transfer at the electrode-solution interface, there is also effective impedance related to diffusion-controlled mass transfer. For example, the situation where Ox is reduced to R can be considered as three steps:

1. Mass transport (by diffusion) of Ox from the bulk electrolyte to the electrode surface,

2. Charge transfer reaction (kinetic control) that converts Ox to its reduced form R,3. Mass transport of R from the electrode surface into the bulk electrolyte.

The overall impedances related to these processes can be represented as a linear combination of a resistor Rs and capacitor Cs. It can be shown that the charge transfer resistance Rct can be separated from the mass transport processes to give the angular frequency (ω = 2πf) dependencies for Rs and Cs as:

; For a planar diffusion front over an electrode of surface area A, σ is given by:

where DOX and DR are the diffusion coefficients for the electroactive species of bulk fluid concentration [OX]b and [R]b, respectively. The impedance Z of the linear combination of Rs and Cs is

which takes the form of a conventional resistance element Rct in series with a frequency-dependent element known as the Warburg impedance ZW:

I = IOx = nF [Ox]skOx

I = IRed = nF [Red]skRed

Rct

=RT

nFIo

RS = Rct +

!1/2CS =

1

!1/2

Z = RS +1

j!Cs= Rct +

!1/2+

j!1/2

ZW = Re(ZW ) + Im(ZW ) =

!1/2 j

!1/2

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The Warburg impedance possesses equal real and imaginary components, and so takes the form of a constant phase element, characterised by a constant phase angle of 45°. The magnitude of ZW is:

(4)The relative values of ZW and Rct provide an indication of the balance between mass transport control and charge transfer kinetics of an electrode reaction. Unless the charge transfer process exhibits very slow kinetics, at high frequencies the Warburg impedance will be negligible compared with the charge transfer resistance, whereas at very low frequencies it will have the dominating influence on an electrode reaction. The total impedance of an electrochemical cell will include contributions from the counter electrode as well as the working electrode. In order to be able to focus on the reaction occurring at the working electrode, the impedance of the counter electrode is reduced to insignificance by making its surface area as large as possible. The impedance of the working electrode should also include the effective capacitance Cdl of the electrical double layer at its surface, which will appear as an element in parallel with the charge transfer resistance and Warburg impedance. The bulk electrolyte between the two electrodes will appear as a series resistance RΩ, and this should also be considered. The geometrical capacitance between the working and counter electrode will be negligibly small, and the reactance it presents will be short-circuited by the bulk electrolyte resistance. The overall equivalent circuit for a simple reversible reaction at a working electrode system can therefore take the form shown in Fig. 1:

! Fig. 1: Equivalent circuit that includes a charge transfer resistance Rct that controls the kinetics of a simple reversible electrode reaction, together with the Warburg impedance Zw that controls the mass transport. The resistance RΩ of the bulk electrolyte and the capacitance Cdl of the electrical double layer at the electrode are also included.

At low frequencies, the real and imaginary impedance components of this equivalent circuit approach the limiting values:

Eliminating the factor σ/ω½ gives the relationship:

(5)

|ZW | =r

!1/2

2+

!1/2

2=

r2

!

RΩ

Rct

Cdl

WZW

Re(Z) = R +Rct +

!1/2

Im(Z) =

!1/2+ 22Cdl

Im(Z) = Re(Z) (R +Rct 22Cdl)

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A plot of –Im(I) versus Re(Z) should thus take the form of a straight line of unit slope, intercepting the Re(Z) axis at (RΩ + Rct – 2σ2Cdl). At very high frequencies, where the Warburg impedance becomes negligible, the real and imaginary impedance components of the equivalent circuit shown in Fig. 1 approach the limiting values:

; Eliminating the frequency ω from these two equations gives:

(6)In the high frequency range, a plot of –Im(Z) versus Re(Z) should thus take the form of a semicircle of radius Rct/2, centred on the Re(Z) axis at (RΩ + Rct/2). A Nyquist plot that incorporates equations (5) and (6) to describe the frequency variation of the impedance of the equivalent circuit of Fig.1 is shown in Fig.2:

! Fig. 2: Nyquist plot for the equivalent circuit of an electrode reaction shown in Fig.1. Mass transfer control operates at low frequencies, and kinetic control occurs at high frequencies.

In Fig. 2 the frequency range of measurement of Re(Z) and Im(Z) would typically extend from 10–3 Hz to 1 MHz, to show the kinetically controlled region of the electrode reaction (the semicircle) and the mass transport (diffusion) controlled region represented by the straight line of unit slope.

References [1] J. Zhu, Z. Zhu, Z. Lai, R. Wang, X. Guo, X. Wu, G. Zhang, Z. Zhang, Y. Wang, and Z. Chen, “Planar amperometric glucose sensor based on glucose oxidase immobilized by chitosan film on prussian blue layer,” Sensors, vol. 2, no. 4, pp. 127–136, 2002.

[2] E. Katz and I. Willner, “Probing biomolecular interactions at conductive and semiconductive surfaces by impedance spectroscopy: Routes to impedimetric immunosensors, DNA-Sensors, and enzyme biosensors,” Electroanalysis, vol. 15, no. 11, pp. 913–947, 2003.

Re(Z) = R +Rct

1 + !2C2dlR

2ct

Im(Z) =!CdlR2

ct

1 + !2C2dlR

2ct

Re(Z)R Rct

2

2

+ (Im(Z))2 =

Rct

2

2

Re(Z)

ω = ∞

‒Im

(Z)

45o

RΩ Rct+RΩ

ω = 1/RctCdl

ω

(Rct + RΩ ‒ 2σ 2Cdl)

KineticControl

Mass TransportControl