introduction to electrochemical biosensors...biosensors and instrumentation stewart smith beijing...

Stewart SmithBiosensors and InstrumentationBeijing University of Posts and Telecommunications 2019

Introduction to Electrochemical Biosensors

Lecture 5

1


Summary• Introduction to Electrochemical

Biosensors ‣ Potentiometric sensors

‣ Amperometric sensors

‣ Oxygen sensing (not a biosensor)

‣ Microelectrodes

‣ Glucose sensing

2


Electrochemical Biosensor

Signal

SampleA/D Converter

TransducerAnalyte

Immobilised Biological Receptor

An Integrated, Biological Receptor-Transducer, Device

3


Electrochemical Sensors• Three main types of sensor ‣ Potentiometric:

- Use Ion Selective Electrodes to determine the concentration of chosen ions.

‣ Amperometric: - Measure current resulting from redox reactions.

‣ Conductometric - Measure changes in ionic composition resulting

from an enzyme reaction.

4


Reference (Counter) Electrode

E Control or

Measurement

I

Electrochemical Sensors• Potentiometric: Measure equilibrium E (I = 0)

• Amperometric: Control E, measure I

Working (Indicator, Detector) Electrode

5



E Control or

Measurement

I




6


Potentiometric Sensor• pH - Common potentio-

metric measurement

• Glass pH Electrode measures concentration of Hydronium (H+) ions

• This, and other potentio-metric sensors require a stable reference electrode

E+ –

Porous FritGlass H+

Membrane

SensingElectrode (Ag-AgCl)

ReferenceElectrode (Ag-AgCl)

7


Potentiometric Biosensor

E

• Potentiometric urea sensor

• Two pH sensors: ‣ Sensing electrode coated

with urease enzyme ‣ Bare reference sensor

• Local pH change at urease coated electrode

• pH (potential) difference between electrodes proportional to urea conc.

Reference pH sensor

Urease

CO(NH2)2 + 3H2Ourease

2NH +4 + HCO –

3 + OH–

8


Potentiometric Biosensors• 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

9



E Control or

Measurement

I




10


Amperometric Oxygen Sensor

©Gary Christian, Analytical Chemistry, 6th Ed. (Wiley)

• Clark Oxygen Electrode

• Oxygen reduced at Pt cathode ring

• Anode coil reference electrode: Ag-AgCl

• Teflon membrane allows O2 diffusion

• Current will depend on reaction rate and pO2

11


Clark Oxygen Sensor

Dissolved Oxygen

analyte solution electrolyte

O2

O2

O2

O2 permeable membrane (must stir solution to avoid

‘diffusion barrier’)Pt electrode

–0.6 to –0.7V

At the Pt cathode:

At the Ag-AgCl anode:

O2 + 4e− + H2O ⟶ 4OH−

Ag + Cl− ⟶ AgCl + e−

id - measured current F - Faraday's constant Pm - permeability of O2

A - electrode area [O2] - oxygen concentration b - thickness of the membrane

id = 4FPmA[O2]b

12


Electrode Surface Reactions

Flux of electrons

Flux of Ox

Flux of R

Flux Balance at the Electrode Surface

Planar Diffusion

Stirring the solution reduces ‘Diffusion

Barrier’ in Mass Transfer.

13


Hemispherical Diffusion

Electrode Surface Reactions

Flux of electrons

Flux of Ox

Flux of R

Flux Balance at the Electrode Surface

Microelectrodes in miniaturised sensors do not require stirring

13


Clark Oxygen Sensor• Wide range of applications:‣ Environmental studies (e.g. O2-levels in natural waters)‣ Sewage treatment (monitoring bacterial treatment).‣ Alcohol production (O2-levels in fermenters need to be

continuously monitored and controlled)‣ Similar uses in other industrial bio-reactors‣ Medicine (invasive and non-invasive monitoring)

• Biosensor Applications - for example, when combined with an immobilised enzyme?

14


Amperometric Glucose Sensor

• Glucose-oxidaze (GOx/GOD) enzyme reaction

• How to measure this?• Measure production of

hydrogen peroxide?• Electrochemical detection:

H2O2 ⟶ O2 + 2H+ + 2e−

• Plateau current depends on H2O2 conc. which depends on glucose conc.

0

Anodic

Cathodic

+i

-i

E+ 0.6 V

15


“First Generation” Sensor Issues

• H2O2 detection may be confounded with interferant compounds such as uric acid.

• Use of oxidases means that oxygen is required, so the reaction is dependent on this concentration.

• How have these issues been addressed?

16


Possible Solutions

Oute

r Mem

bran

e

Immo

bilise

d GO

x Lay

er

Inner

Mem

bran

e

Elec

trode

Glucose, Oxygen

H2O2

Interferants

• Use of selectively permeable membranes

• Outer controls O2 and glucose flux

• Inner prevents transport of interferants

• Other solutions for this involve catalysing H2O2 reactions to lower potential

17


Reactions in GOx • Flavin adenine dinucleotide

(FAD) at the heart of GOx

• Redox active site:

‣ FAD - Oxidised form

‣ FADH2 - Reduced form

• Interacting directly with the FAD could replace oxygen

• Difficult because it is buried away within the molecule

GOx(FAD) + glucose→ GOx(FADH2) + glucono-lactone

GOx(FADH2) + O2 → GOx(FAD) + H2O2

FAD

18


“Second Generation” Sensors

• Oxygen is replaced by mediator (Med):

• Mediator is oxidised at electrode to transfer electrons:

• Common mediators include ferrocene, ferricyanide, quinones and phenothiazone

GOx(FAD) + glucose + H2O ⟶ GOx(FADH2) + gluconolactoneGOx(FADH2) + 2Med(Ox) ⟶ GOx(FAD) + 2Med(Red) + 2H+

nMed(Red) ⟶ nMed(Ox) − ne−

19


Mediated Electron Transfer

Electrode biased at Oxidising Voltage

e–

i Reduced Enzyme

Oxidised Enzyme

Substrate (Analyte)

Product

Ferrocyanide

Ferricyanide

20


Mediated Electron Transfer

e–

i GOx(FADH2)(Red)

GOx(FAD)(Ox)

Glucose

Glucono-lactone

Ferrocyanide (Red)

Ferricyanide (Ox)

Electrode biased at Oxidising Voltage

21


Mediator Characteristics• Mediator allows the biosensor ‣ to be independent of oxygen concentration

‣ to operate at lower potentials

• To be effective, the mediator should ‣ react rapidly and preferentially with the enzyme

and at the electrode

‣ be highly soluble and diffuse quickly

‣ be non-toxic and chemically stable

22


Direct Electron Transfer

Electrode biased at Oxidizing Voltage

e– Substrate (Analyte)

Product

Transduced Current i

23



• Cellobiose dehydrogenase (CDH) enzyme

• Internal charge transfer from FAD to cytochromeImage from: R. Ludwig, R. Ortiz, C. Schulz, W. Harreither, C. Sygmund, and Lo Gorton, “Cellobiose dehydrogenase modified electrodes: advances by materials science and biochemical engineering,” Analytical and Bioanalytical Chemistry, vol. 405, no. 11, pp. 3637–3658, Jan. 2013.

24



Electrode biased at Reducing Voltage

e– Substrate (Analyte)

Product

Oxidase

H2O2

O2H2O

Peroxidase

Direct, non-mediated, electrical contact of two EnzymesTransduced

Current i

25


Further Developments

• Direct electrical “wiring” of enzymes to electrodes?

26


Implantable Glucose Sensors

• Holy grail for diabetes control

• Needs to work for days/weeks/months

• Typically they fail within days

• Subject of huge research effort

bovine serum albumin (BSA), and glutaraldehyde (50%) were obtainedfrom Fisher Scientific (Pittsburgh, PA). !0.125 mm Teflon-coveredplatinum-iridium (9:1 in weight) and silver wires were obtained fromWorld Precision Instruments, Inc. (Sarasota, FL). High-temperature miniround glue sticks were supplied by Adhesive Tech (Hampton, NH).Hematoxylin and eosin were obtained from Fisher Scientific (Pittsburgh,PA), and trichrome stains were purchased from Newcomer Supply(Middleton, WI).

Synthesis of Poly(HEMA-DHPMA-VP-EGDMA) HydrogelSeries. Various molar ratios of HEMA and DHPMA were weighedand mixed in 20 mL glass vials. VP (9 mol %), 1 mol % EGDMA,and 1.5 mol % initiator Benacure 1173 were added to the solutions(Table 1). The solutions were diluted 1:1 with distilled water. Afterdeoxygenation with argon gas, the solutions were injected intohandmade glass cells measuring 55 mm × 25 mm × 1 mm. The sidesof the cells were secured with a glue gun. The solution filled cellswere placed under a UV lamp (Spectronics Corporation, Westbury,NY) at a wavelength of 254 nm for 100 min. Argon gas was used topurge the system in order to minimize oxygen inhibition of the reaction.Transparent uniform sheets were obtained, and these xerogels wereswollen in distilled water to yield hydrogels. The water was refreshedevery day for one week to remove any unreacted monomers.

The hydrogels were freeze-dried for various studies using a VirtisFreezemobile 12XL freeze-dryer (The Virtis Company Inc., Gardiner,NY).

Scanning Electron Microscopy (SEM). The hydrogel samplesobtained by the above method were cut and freeze-dried to determinethe inner morphology. The morphology of the freeze-dried gels werestudied via a Hitachi scanning electron microscope S-800 (Hitachi HighTechnologies, Pleasanton, CA) with 20 nm gold coating by a HummerXsputter coater (Anatech, Ltd., Springfield, VA). The working distancebetween the sample and electron (WD) was 5 mm with 25.0 kV.

Sorption Experiments. Hydrogel samples were cut with a 12 mmdiameter cork bore to obtain uniform shapes. All sheets were freeze-dried for one day and stored in a desiccator at room temperature. Thesorption behavior of hydrogels was monitored by detecting the increasein mass of the samples at different time intervals by a Sartorius BP211Dbalance ((0.01 mg, Sartorius Corporation, Edgewood, NY). In a typicalsorption experiment, a preweighed dry gel sheet was immersed intowater at 24 ( 1 °C in a Fisher Scientific Isotemp water bath (Pittsburgh,PA). At prescribed time intervals, the hydrogel was taken out of solutionand weighed after wiping off the excess water from the surface withKimwipe paper (Kimberly Clark Professional). The sorption degree,SD, of hydrogels was defined as follows:

SD% )Wt -Wd

Wd× 100 )

Mt

Wd× 100 (1)

where Wd is the weight of the dry gel, Wt is the weight of wet hydrogelat each time interval, and Mt is the gain in the weight of the dry gel attime t.

The equilibrium water content (EWC) of the hydrogel was deter-mined using the following equation:

EWC% )W∞-Wd

W∞× 100 )

M∞

W∞× 100 (2)

where W∞ and M∞ are the weight and weight gain of the swollenhydrogel at equilibrium, respectively.

Differential Scanning Calorimetry (DSC). All calorimetric datawere obtained via a TA Instruments 2920 differential scanningcalorimeter (DSC, TA Instruments, New Castle, DE). Nitrogen gas waspassed through the instrument at a flow rate of 70 mL/min. Beforemeasurement, the DSC was calibrated from -100 to 250 °C at a heatingrate of 5 °C/min with an indium standard. All the sample masses forDSC ranged from 4 to 10 mg.

For the determination of glass-transition temperature, samples werefreeze-dried for one day and kept in a desiccator. The samples werequickly weighed, sealed in aluminum pans, and immediately scannedfrom 30 to 200 °C. Samples were cooled and rescanned; the glasstransition was determined from the second heating cycle in order tominimize any aging effects. The reported glass-transition temperaturewas determined as the midpoint at half heat flow or heat capacity.

To measure the water structure of equilibrium swollen hydrogels,first they were cut and weighed accurately after wiping off the surfacewater with Kimwipe paper and immediately sealed in Al pans. TheDSC curves were measured on heating from -100 to 20 °C with aheating rate 5 °C/min. Endotherm areas were determined.

Sensor Preparation. Hydrogel-coated glucose sensors used for thisstudy were based on the coil-type implantable sensor previouslydescribed.20,21 The detail fabrication method of Pt/GOx and Pt/GOx/epoxy-polyurethane-based sensors can be found in those papers. Thecoating process was achieved by applying a ∼0.4 µL of 50%hydrogel–solution to either the enzyme layer or the epoxy-polyurethanelayer, using a 10 µL micropipette, and then curing under an argonatmosphere and a 254 nm UV light for 100 min. Membrane surfaceswere observed and photographed using an Olympus BX41 microscope(Quantitative Imaging Co., Canada). The cured hydrogel coating wastransparent and firmly attached to the sensor as shown in Figure 1.

In Vitro Evaluation Method. Electrochemical measurements wereperformed with Apollo 4000 4-channel potentiostats (World PrecisionInstruments, Inc., Sarasota, FL). Newly prepared glucose biosensorswere conditioned in a 5 mM glucose/PBS (NaCl 8.76 g, KH2PO4 3.53 g,Na2HPO4 3.40 g, benzoic acid 2 g, water 10 000 mL) (ionic strength) 0.16) for at least 2 h and then continuously polarized at +0.7 V vsAg/AgCl until a stable background current was reached. The responsewas determined as the time needed for the sensor’s current to reach90% of (T90%, min), the maximum current when the glucose concentra-tion changed from 5 mM to 15 mM. The sensitivity (S) was determinedusing a two-point method and can be calculated by

S(nA/mM) )I15mM - I5mM

10(3)

where I15mM and I5mM represent the sensor response currents obtainedin 15 mM and 5 mM glucose solutions, respectively. The sensors forlong-term observation were stored in PBS. Calibration plots wereobtained by stepwise adding 100 mM glucose solutions to 8.0 mL PBS.All measurements were performed at 24 ( 1 °C.

In Vivo Biocompatibility Studies. The epoxy-PU-coated Silastictubing (Dow Corning, Midland, MI) (! ) ∼ 4 mm, length ) ∼50

Table 1. Feed Compositions of Synthesized Xerogels

mol % HEMA DHPMA VP EGDMA

DHPMA0 90 0 9 1DHPMA20 70 20 9 1DHPMA40 50 40 9 1DHPMA60 30 60 9 1DHPMA80 10 80 9 1DHPMA90 0 90 9 1

Figure 1. Sensor photo showing a thin layer of hydrogel coating.

562 Biomacromolecules, Vol. 9, No. 2, 2008 Wang et al.

27


Summary - Glucose Sensors• 1st Generation glucose sensor: ‣ Measure enyzme reaction products with

amperometric sensor

• 2nd Generation: ‣ Use a mediator to transfer electrons to the

sensing electrode

• Future generations ‣ Direct “wiring” of enzymes to electrodes

28


Conductometric Biosensors• Detect changes in Electrical Conductivity

resulting from an Enzyme Reaction.Sources of Conductivity Change Enyzmes

Generation of ion groups Amidases

Separation of different charges Dehydrogenases & decarboxylases

Ion migration (proton conduction) Esterases

Change in association of ion groups Kinases

Change in size of charged groups Phosphatases & sulphatases

29


Example Reactions

Urea

D amino acid + O2 + H2ODAAO

↵ keto acid– + H2O2 + NH +4

CO(NH2)2 + 3H2Ourease

2NH +4 + HCO –

3 + OH–

30


Conductometric Biosensors• Interdigitated electrode

design

‣ Screen printing of platinum or silver-palladium materials?

‣ Vacuum deposition of metals and photolithography

‣ Substrate of alumina or glass

‣ Immobilised enzyme covalently bound to electro-inactive protein (e.g., albumin)

31


Conductometric Instrumentation

ZsensorZ1

Z2 Z3

Diff.Amp

+

−

32


Conductometric Instrumentation

−

+

Vout

RfGsensor

Vin

Vout = VinRfGsensor

33


Impedimetric Biosensors

The electron transfer resistance can be translated into theexchange current under equilibrium, Io, Equation 7, andthen the heterogeneous electron transfer rate constant, ket,can be evaluated, Equation 8, where R! 8.31 J mol"1 K"1 isthe gas constant, T is the temperature (K), F! 9.65# 104 Cequiv"1 is the Faraday constant,A is theelectrodearea (cm"2),[S] corresponds to the bulk concentration of the redox probe(mol cm"3), and n is the number of electrons transferred permolecule of the redox probe.

Ret!RT (nFIo)"1 (7)

Io! nFAket[S] (8)

In the absence of any redox label in the electrolyte solution,only non-Faradaic impedance is operative. The electrontransfer characterizing parameters, Ret and ZW, becomeinfinite, and the equivalent circuit can be simplified, asshown in Scheme 1E. The variable component in the circuitis presented by Cmod, and it affects the imaginary part of theimpedance,Zim, Equation 9, considering only the dense partof the double-charged layer.

Zim ! 1! CAu $ Cmod% & %9&

The experimental results can be analyzed graphically [38,39] (in the case of Faradaic impedance spectra usually by theuseof theNyquist coordinates:Zim vs.Zre) in the frameof thetheoretical model [40, 41]. A computer fitting of theexperimental data to a theoretical model represented byan equivalent electronic circuit is usually performed. All

electronic characteristics of the equivalent circuit and thecorresponding physical parameters of the real electrochem-ical system can be extracted from such analysis. Since thevariable parameters of the system represent the functions ofthe modifying layer and its composition, they can be used toquantitatively characterize the layer. Analysis of the Zre(!)and Zim(!) values observed at different frequencies allowsthe calculation of the following important parameters: a) thedouble-layer capacitance, Cdl and its variable component,Cmod; b) the electron transfer resistance,Ret, and its variablecomponent, Rmod; c) the electron transfer rate constant, ket,for the applied redox probe derived from the electrontransfer resistance, Ret. Thus, the impedance spectroscopyrepresents not only a suitable transduction technique tofollow the interfacial interactions of biomolecules, but it alsoprovides a very powerful method for the characterization ofthe structural features of the sensing interface and forexplaining mechanisms of chemical processes occurring atthe electrode/solution interfaces [45].

3. Immunosensors Based on ImpedanceSpectroscopy

Among the biomaterial-based sensing devices immunosen-sors are anticipated to be an important class of sensingsystems in clinical diagnosis, food quality control, environ-mental analysis, detection of pathogens or toxins, and eventhe detection of explosives or drugs [46 ± 48]. Severalmethodologies for the electrochemical transduction of theformation of antigen-antibody complexes on the electrodesurface were reported [24, 49]. Impedance spectroscopy(including Faradaic impedance in the presence of a redoxprobe and non-Faradaic ± capacitance measurements)represents an important electronic transduction means ofantigen-antibody binding interactions on electronic ele-ments.

3.1. Immunosensors Based on In-Plane ImpedanceMeasurements Between Electrodes

In-plane impedance measurements were performed foranalyzing antigen-antibody recognition processes. The de-tection path is based on the organization of two metallicconductive electrodes on an electrical insulator leaving anon-conductive gap between the electrodes, Scheme 3A.The gap was modified with a biomaterial capable of specificbinding of a complementary unit, e.g., the gap was modifiedwith antigen molecules for binding of the complementaryantibody molecules, or an antibody was linked to the gap tobind the respective antigen molecules. Affinity binding ofthe respective complementary component results in thechange of the electrical properties of the gap, thus affectingthe impedance between the electrodes. The impedancechanges can bemeasured as variation of conductivity,"gap!R"1

gap, between the electrodes, that represents changes in thereal part, Zre, of the complex impedance [50] or variation of

Scheme 3. A) Immunosensor for the in-plane impedance meas-urements between the conductive electrodes. B) Interdigitatedelectrode for the in-plane impedance immunosensing.

916 E. Katz, I. Willner

Electroanalysis 2003, 15, No. 11

• Capacitive sensing

• Small AC signal applied

• No DC component

• Measure changes in dielectric properties

• Requires sensitive measurement

34


Impedimetric Biosensors

• Interdigitated micro-electrode design

• Various sensor designs measure changes in R or C

35


Electrochemical Impedance Spectroscopy

• Theory

• Techniques

• Equivalent circuits and impedance plane plots

• Applications

36


Electrochemical Impedance Spectroscopy (EIS)

• Butler-Vollmer equation defines redox current

• Small (5-10 mV) AC signal centred on Eo

• Ignore higher order effects for linear output

• Effects from charge transfer (kinetics) and mass transport

Ox + ne– ↔ Red

Cur

rent

(A

)Voltage (V)

E1/2 = Eo

I = IonF

RT(E Eo)

37


I

EE0

I0

Impedance Spectroscopy 1• Apply small AC

signal: ‣ E(ω) = Eo + ΔE

sin(ωt)

• Measure response: ‣ I(ω) = I0 + ΔI

sin(ωt+φ)

• Frequency, ω = 2π f

38


I

EE0

I0

Impedance Spectroscopy 1• Apply small AC

signal: ‣ E(ω) = Eo + ΔE

sin(ωt)

• Measure response: ‣ I(ω) = I0 + ΔI

sin(ωt+φ)

• Frequency, ω = 2π f

EoΔE

I0

ΔI

38


Impedance Spectroscopy 2• If ΔE is small

then the response should be linear

• Impedance is then:

• Complex value:

E0

ΔE

I0

ΔIZ(!) =E(!)

I(!)

Z(!) = |Z(!)|ej(!)

39


Electrochemical Impedance Spectroscopy (EIS)

• Butler-Volmer equation defines redox current

• Small (5-10 mV) AC signal centred on E0

• Ignore higher order effects for linear output

• Effects from charge transfer (kinetics) and mass transport

Ox + ne– ↔ Red

Cur

rent

(A

)Voltage (V)

E1/2 = Eo

I = I0nF

RT(E E0)

40


Contributions to EIS response

• Three main components of the impedance from a Faradaic reaction: ‣ Diffusion of Ox to electrode surface from bulk

‣ Reaction kinetics of Ox + ne– ↔ Red

‣ Diffusion of Red to bulk solution

• Charge transfer resistance Rct and “Warburg Impedance” ZW

ne–Ox

Red

41


Other Contributions to Electrochemical Impedance

• Counter electrode is large so that it does not limit the current (Rct(CE) ≪ Rct(WE))

• Significant capacitance from the electrical double layer (which we’ll get to)

• Cdl appears in parallel with Rct, ZW

• Final component is related to the solution conductivity: RΩ

42


Electrical Double Layer Capacitance

• Cdl made up of two parts, compact “Stern layer” and the diffuse layer

• Diffuse layer width decreases with solution concentration, increasing Cd

+−

−

−

−

−

+

+

+

+

−

−−

−

+

+

+

+

−

−

−

−

+

+

+

+

Elec

trode

Sur

face

HelmholtzLayer

Diffuse Layer

CH Cd

1

Cdl

=1

CH

+1

Cd

43


Randles Equivalent Circuit

• At low frequencies the limiting values of Z:

• Eliminate to leave:

σω1/2

R

R

C

ct

dl

Ω

WZw

Re(Z) = RΩ + Rct + σω1/2 −Im(Z) = σ

ω1/2 + 2σ2Cdl

−Im(Z) = Re(Z) − (RΩ + Rct + 2σ2Cdl)

44


Randles Equivalent Circuit

• At high frequencies the limiting values of Z:

• Eliminate to leave:

ω

R

R

C

ct

dl

Ω

WZw

Re(Z) = RΩ + Rct

1 + ω2C2dlR2ct

−Im(Z) = ωCdlR2ct

1 + ω2C2dlR2ct

(Re(Z) − RΩ − Rct

2 )2

+ (−Im(Z))2 = ( Rct

2 )2

45


Complex Impedance• Polar co-ordinates: Z(ω) = |Z(ω)|e jϕ(ω)

• |Z| magnitude, ϕ phase shift

• Cartesian: Z(ω) = Zr(ω) + jZj(ω)

• Zr real part, Zj imaginary part

• |Z| = √(Zr2 + Zj2), ϕ = tan–1(Zr / Zj)

• Zr = |Z|cos(ϕ), Zj = |Z|sin(ϕ)

46


Impedance Plane Plots• Plot of the complex impedance plane

• Related to Argand, Nyquist and Cole-Cole plots

• Real part of Z on X-axis

• Imaginary part of Z on Y-axis

• Each point on the curve is a particular frequency ω

47


Impedance Plane Plots

Increasing ω

−Im Z

Re ZR

Z = R − j/ωC

C

R

48



• Double layer capacitance Cdl

• 20-60 μF cm–2 • Charge transfer

resistance • Faradaic reaction

−Im

Z

Re ZRct0

ω = 0ω

ω = ∞

ϕ|Z|

RCdl ctω = 1/RctCdl

49



• Includes solution resistance RΩ

• Series resistance shifts response along Re Z axis

−Im

Z

Re ZRΩ + Rct0

ω = 0ωω = ∞

RΩ

Cdl

RctRΩ

50



• At low frequencies:

• Straight line, unit gradient, intersects axis at:

R

R

C

ct

dl

Ω

WZw

Re(Z)

ω = ∞

‒Im

(Z)

45o

RΩ Rct+RΩ

ω = 1/RctCdl

ω

(Rct + RΩ ‒ 2σ 2Cdl)

KineticControl

Mass TransportControl

−Im(Z) = Re(Z) − (RΩ + Rct + 2σ2Cdl)

Re(Z) = RΩ + Rct + 2σ2Cdl

51



• At high frequencies:

• (Semi)circle with radius Rct /2, centre:

R

R

C

ct

dl

Ω

WZw

Re(Z)

ω = ∞

‒Im

(Z)

45o

RΩ Rct+RΩ

ω = 1/RctCdl

ω

(Rct + RΩ ‒ 2σ 2Cdl)

KineticControl

Mass TransportControl

(Re(Z) − RΩ − Rct

2 )2

+ (−Im(Z))2 = ( Rct

2 )2

Im(Z) = 0, Re(Z) = RΩ + Rct

2

52


Real Data

Organic LED

Charge transport in organic semiconductors

RCdl ct

“Rct”

At maximum Im Z, f = 1.1 kHz

53


Bode Plot 54


Experimental Setup • 2, 3 or 4 terminal measurements • Potentiostatic mode ‣ DC potential with sinusoidal voltage ‣ Most common setup

• Galvanostatic mode ‣ DC with added sinusoidal current ‣ Electrodeposition measurements

55


Electrochemical Cell

WE

RE

CE• RE: Reference

Electrode

• CE: Counter (or auxiliary) Electrode

• WE: Working Electrode

56


Electrochemical Cell

• RE: Reference Electrode

• CE: Counter (or auxiliary) Electrode

• WE: Working Electrode

56


Biosensor Applications

• Review paper: E. Katz and I. Willner, Electroanalysis, vol. 15, 2003

• EIS is ideal for the measurement of changes to a surface caused by attachment of biomolecules.

57



• Functionalised microelectrodes

• Attachment changes EIS response

•Measurement of changes in Rct or Cdl

chemical flow cells were introduced to compare theimpedance spectra in the two detection channels: with andwithout the analyte sample [63, 81]. Experimental [81] andtheoretical [63] details on the use of differential impedancespectroscopy for immunosensing were discussed.Another point that has been addressed is the often-

observed poor reversibility of the immunological reactions.The binding between the antigen and the antibody is usuallyvery strong. By applying acidic media or high concentrationof salts or urea, the proteins associated with the antigen-antibody complex are unfolded, leading to separation ofantigen and antibody. These severe conditions affect theprotein that acts as the receptor component of the sensorand the reactivation of the sensing interface after separationof the antigen-antibody complex is often impossible.Even inthe case, if the receptor component is not a protein and it hashigh chemical stability (e.g., dinitrophenyl antigen, DNP-antigen), the structure of the monolayer could be perturbedand the thiol anchor groups could be dissociated from thesurface upon the conditions required for the affinitycomplex dissociation. Photoisomerizable antigenmoleculeswere suggested as a means to tailor reversible and reusableimmunosensor electrodes [75, 76, 82, 83], Scheme 7A. Bythis approach the antigen inphotoisomer stateAenables thesensing of the antibody by the electronic transduction of the

Fig. 2. Non-Faradaic impedance analysis of the foot-and-mouth-disease antibody on an electrode surface functionalized with therespective antigen (3). A) The absolute impedance value, !Z ! , as a function of the applied frequency: a) Prior to the antibody binding,b) After the formation of the saturated antigen/antibody complex film. B) The imaginary part of the impedance, Zim, as a function of theapplied frequency: a) Prior to the antibody binding. b) After the formation of the saturated antigen/antibody complex film. C) Theelectrode capacitance changes upon the immunosensing: a) Using the antibody at a concentration of 1.74 mg mL" 1. b) Using theantibody at a concentration of 17.4 mg mL" 1. c) Using a non-specific antibody at a concentration of 17.5 mg mL" 1. (Reproduced from[66] with permission).

Fig. 3. Faradaic impedance spectra recorded upon sensing offoot-and-mouth-disease antibodies at a Au electrode functional-ized with a self-assembled monolayer of the thiolated polypeptide(3) corresponding to the foot-and-mouth-disease antigen: a) BareAu electrode. b) (3)-monolayer functionalized Au electrode priorto the reaction with the antibody. c) (3)-monolayer functionalizedAu electrode after the reaction with the antibody. The data wererecorded in the presence of [Fe(CN)6]3" /4" , 2 mM. (Reproducedfrom [65] with permission).

921Impedimetric Immunosensors, DNA-Sensors, and Enzyme Biosensors

Electroanalysis 2003, 15, No. 11

(a) Bare electrode(b) Functionalised electrode(c) With antibody attachment

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• EIS biosensor development at Edinburgh

• Gold electrodes with bound immunosensing molecules (antibodies)

• Binding of antigen increases Rct[1] I. Ciani, et al., “Development of immunosensors for direct detection of three wound infection biomarkers at point of care using electrochemical impedance spectroscopy,” Biosens Bioelectron, vol. 31, no. 1, pp. 413–418, Jan. 2012.

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• Lectin (Con A) used to selectively bind cancer cells

[1] Y. Hu, P. Zuo, and B.-C. Ye, “Label-free electrochemical impedance spectroscopy biosensor for direct detection of cancer cells based on the interaction between carbohydrate and lectin,” Biosens Bioelectron, vol. 43, pp. 79–83, May 2013.

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EIS - Microelectrodes

• Not diffusion limited at low f

• Modified Randles equivalent circuit

• ZW ⇒ 0 as f ⇒ 0 • Rnl - non-linear

R

R

C

ct

dl

Ω

WZw

nlR

−Im

Z

Re ZRΩ + Rct

0

ω = 0ωω = ∞

RΩ

RΩ + Rct + Rnl

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EIS - Microelectrodes• 50µm Pt micro-

electrode • EIS performed in

solution of: ‣ 5mM ferricyanide/

5mM ferrocyanide • DC = 0.19 V • f:1000 Hz to 0.1

Hz!

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EIS Summary

• Electrochemical Impedance Spectroscopy ‣ Measurement basics, small AC signal at Eo

‣ Contributions to impedance and equivalent circuits

‣ Impedance plane plots and physical relevance

‣ Applications in Biosensing

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introduction to electrochemical biosensors...biosensors and instrumentation stewart smith beijing...

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