Integrated electrochemical DNA biosensors for lab-on-a-chip devices

Review

Integrated electrochemical DNA biosensorsfor lab-on-a-chip devices

Analytical devices able to perform accurate and fast automatic DNA detection or sequen-

cing procedures have many potential benefits in the biomedical and environmental fields.

The conversion of biological or biochemical responses into quantifiable optical, mechanical

or electronic signals is achieved by means of biosensors. Most of these transducing

elements can be miniaturized and incorporated into lab-on-a-chip devices, also known as

Micro Total Analysis Systems. The use of multiple DNA biosensors integrated in these

miniaturized laboratories, which perform several analytical operations at the microscale,

has many cost and efficiency advantages. Tiny amounts of reagents and samples are

needed and highly sensitive, fast and parallel assays can be done at low cost. A particular

type of DNA biosensors are the ones used based on electrochemical principles. These

sensors offer several advantages over the popular fluorescence-based detection schemes.

The resulting signal is electrical and can be processed by conventional electronics in a very

cheap and fast manner. Furthermore, the integration and miniaturization of electro-

chemical transducers in a microsystem makes easier its fabrication in front of the most

common currently used detection method. In this review, different electrochemical DNA

biosensors integrated in analytical microfluidic devices are discussed and some early stage

commercial products based on this strategy are presented.

Keywords:

DNA / Electrochemical DNA biosensors / Electrochemistry / Lab-on-a-chip /Micro Total Analysis systems DOI 10.1002/elps.200900319

1 Introduction

One of the analytical applications that have caught more

attention during the last 25 years is DNA detection and

sequencing. This is because it has many potential benefits in

the biomedical and environmental fields, and also in other less

popularly known areas such as forensics or archeology.

Integrating DNA sensors with lab-on-a-chip (LOC) devices

opens the possibilities of developing portable equipments that

would be very useful for point-of-care medical screenings or

any on-the-field analysis. By definition, LOC devices are

microfluidic devices able to perform complete laboratory

assays that can be alternatively referred as Micro Total Analysis

systems when the application is of an analytical nature.

The automation and integration of one or several

laboratory operations in small devices goes back to the early

1990s. It was soon regarded as a means to obtain higher

efficiency, faster analysis time, and lower reagent

consumption, than other existing analysis techniques [1]. In

1998, a more advanced device was presented [2]. A single

closed system with dimensions of a few squared centimeters

and a height of 1 mm was developed in order to prepare and

analyze nanoliter-sized DNA samples by gel electrophoresis,

containing microfabricated fluidic channels, heaters,

temperature sensors, and fluorescence detectors.

The main aim of developing miniaturized and inte-

grated devices is to reduce the assay cost and also the

interaction of the sample with external factors, reducing the

sample contamination and making the analysis cleaner and

less hazardous. However, they also present some other

advantages derived from scaling down chemical processes to

the microscale such as diffusion mixing efficiency

enhancement, improvement of heat transport, decreased

reaction time, and formation of by-products, bringing as a

result higher reaction yields [3].

DNA sample preparation steps, including complex

processes such as the PCR have been demonstrated to be

achievable on the microscale making such LOC devices very

suitable for complete DNA analysis when integrating

detection elements. The PCR is a very desirable step when

only tiny amounts of the DNA sample are available since it is

Monica Mir1,2

Antoni Homs1,2,3

Josep Samitier1,2,3

1Nanobioengineering group,Institute for Bioengineering ofCatalonia, Barcelona, Spain

2Centro de InvestigacionBiomedica en Red enBioingenierıa, Biomateriales yNanomedicina, Spain

3Department of Electronics,University of Barcelona,Barcelona, Spain

Received May 15, 2009Revised July 10, 2009Accepted July 11, 2009

Abbreviations: ALP, alkaline phosphatase; CG, colloidalgold; CMOS, complementary metal-oxide-semiconductor;CNT, carbon nanotubes; FET, field-effect transistors; HRP,

horseradish peroxidase; a-HL, a-hemolysin; IDE,

interdigitated electrode; LOC, lab-on-a-chip; NTFET,

nanotube field-effect transistor; NW, nanowires; SiNW,

silicon nanowires; SWNT, single wall nanotube

Correspondence: Dr. Monica Mir, Intitut de BioEnginyeria deCatalunya Baldiri Reixac, 13, Barcelona 08028, SpainE-mail: [email protected]: 134-9340319702

& 2009 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim www.electrophoresis-journal.com

Electrophoresis 2009, 30, 3386–33973386

able to amplify a single or few copies up to several orders of

magnitude. This technique is based on the application of

repeated thermal cycles to replicate a target DNA on a solution

containing the mediating DNA polymerase enzyme and

primers (short DNA fragments) complementary to the target.

The increasing demand of cheap, highly accurate, fast

and easy to use analytical assays in the fields of molecular

diagnostics, drug discovery, and pharmaceutical screening

boosted the research in this field at the end of the last

century. The LOC concept has achieved a wide popularity

since its first proposal had proven an effective way to achieve

highly sensitive and complex laboratory assays with mini-

mal effort from the analyst, even it can be used by non-

skilled persons. It has also opened the possibilities of

portable analysis for on-the-field applications due to the

highly reduced dimensions of the devices. Furthermore,

their reduced cost and effectivity makes them a suitable

candidate to be used for global public health applications

even in the third world [4]. Whereas optical detection

methods with fluorescent dyes have dominated the DNA

sensor industry and surface plasmon resonance and piezo-

electric techniques can achieve low limits of detection, the

application of electrochemical read out can provide signifi-

cant advantages in LOC devices. Electrochemistry detection

provides high sensitivity, small dimensions, low cost, fast

response, easy signal integration, and compatibility with

microfabrication technology, making this transducer the

best candidate for the integration in microchip devices.

Accordingly to their detection working principle elec-

trochemical DNA biosensors can be classified as ampero-

metric, voltamperometric, potentiometric, or impedimetric.

Generally, biomolecular recognition takes place close to the

surface of an electrode. For this reason, the electrode

shape, dimension, construction material, and surface prop-

erties play a very important role on the performance of

these biosensors. In most of the cases, the complete

electrochemical detection is achieved by a combination

of a working electrode where the sensing takes place and

a counter reference electrode that should maintain a

stable potential not dependent on the particular detection

reaction but only on the media. These factors have to be

taken into deep consideration when miniaturizing this type

of sensors.

Integration of electrochemical DNA biosensors on

analytical microfluidic devices is a promising approach that

has been already developed in research, as well as in some

commercial biomedical products that are discussed through

the article. However, it still presents several challenges that

should be addressed in the fore coming years in order to

achieve their massive use.

2 Amperometric DNA biosensors

These types of sensors are generally based on the measure-

ment of resulting current in oxidation or reduction

processes of electroactive species in a biochemical reaction

[5]. Amperometric detection in biosensors generally involves

the use of enzymes as biomarker molecules. The read-out of

the enzyme needs the interaction with a specific substrate,

which is converted catalytically in a different molecule. The

product obtained is a redox molecule that can be detected

with this technique and with the advantage that the enzyme

catalyses the reaction, generating signal amplification.

Although the majority of amperometric sensors immo-

bilize the bioreceptor on the electrode surface, in the case of

enzymatic labels it is not necessary, while the redox product

generated with the substrate can diffuse to the electrode in

order to be detected. In this case, the DNA probe is attached

on the PDMS channels [6]. The PDMS channels were

oxidized with plasma, in order to activate the surface for

silane coupling. The silane molecules bear thiol groups

where the maleimide labeled capture probes are attached,

where the target labeled with alkaline phosphatase (ALP)

hybridize. Once it is immobilized on the channel, the

substrate of the enzyme is injected and a redox active

molecule is produced which flows until reaching the indium

tin oxide interdigitated electrode (IDE). In this work a limit

of detection of 1 nM was reported. However, reproducibility

was not tested and the plotted results indicate that the 1 nM

signal had minimum differences with the blank, a more

reliable detection limit for their experiments being 10 nM.

This detection configuration provides a large bioreceptor

surface and since these molecules are not immobilized on

the electrode, as commonly in electrochemical biosensors, it

does not present a barrier that hinders access and detection

of the redox molecules. However, this system is difficult to

use in an array format, since each array spot would need a

specific microchannel.

The combination of amperometric sensors with inte-

grated electronic circuitry and components have recently

brought to market biochips with huge electrochemical

genetic arrays [7]. Combimatrix has commercialized a LOC

device having an integrated multielectrode array able to

perform parallel immunoassays to detect viruses and

bacteria in complex biological samples [8]. The chip has

sample processing capabilities and the sensing part is

composed of individually addressable electrodes controlled

by active complementary metal-oxide-semiconductor

(CMOS) circuitry. This company presented and oligonu-

cleotide microarray platform that provides 12 544 platinum

electrodes, having a diameter of 44 mm/cm2 and coated with

a porous reaction layer. A platinum grid surrounding the

electrodes works as a counter electrode. Different oligo-

deoxynucleotide probes were synthesized at each sensing

microelectrode.

The sensing approach is based on the detection of redox

active chemistries proximal to specific microarray electro-

des. Microarray probes are hybridized to biotin-labeled

targets and then the horseradish peroxidase (HRP)–strep-

tavidin conjugate binds to the biotin molecule, causing an

enzymatic oxidation of the electron donor substrate. The

hybridization is detected by the current generated in the

electro-reduction of the HRP. It is important to notice that

Electrophoresis 2009, 30, 3386–3397 Microfluidics and Miniaturization 3387


this complete platform obtained lower detection limits

(0.75 pM) than its fluorescent detection counter part

(1.5 pM), characterizing the gene expression assay [8]. This

platform has been used to develop nucleic acid assays for

highly accurate genotyping of a variety of common and bio-

thread pathogens.

2.1 Amperometric sandwich

One of most common electrochemical biosensor platform

that has been adapted to be integrated in microfluidic

systems is the sandwich configuration that permits label

free target detection. In the sandwich system two capture

probes are involved, one immobilized on the electrode

surface, where the target hybridize, and a second tagged

probe affine to the non-hybridized segment of the target

(Fig. 1).

In the work reported by Liao et al. [9, 10] seven different

thiolated capture probes from different bacterial species

were self-assembled on screen-printed gold electrodes. A

DNA sequence of 16S rRNA was amplified with PCR. This

technique permits the in vitro cloning of a specific DNA

sequence as well as the labeling of the amplified sequence.

The amplified sequence was hybridized on the capture

probes of the differently functionalized screen-printed

electrochemical cells. The signaling HRP marked serves to

label the target upon hybridization. The electroreduction

current is generated by hydrogen peroxide mediated by

tetramethylbenzidine), obtaining a low detection limit of

12 pM. The microchip design was very simple and it

was composed of 16 individual electrochemical cells

connected to the same microchannel. This channel is inte-

grated to an external macro system for separation and

concentration of pathogens in urine specimens. Conse-

quently, reported pre-treatment is still far to be an integrated

micro-laboratory.

In other research works, site addressing by magnets has

been used in microfluidic devices in order to concentrate

and attach on the sensing electrodes surface oligonucleotide

targets. The research group of Prof. Mascini reported an

enzyme-linked sandwich on magnetic beads for PCR-

amplified gene sequence of Cor a 1.04 allergen detection

[11]. The sensor platform developed starts with a streptavi-

din-coated magnetic beads. Owing to the strong streptavi-

din–biotin interaction, this coated particles allow an easily

and stable biotinylated capture probe immobilization. The

PCR-amplified oligonucleotide target was incubated with

the functionalized particles and this hybridization event was

enlightened by means of a second biotinylated probe that

interact with the non-hybridized part of the target sequence.

This third oligonucleotide interacts with streptavidin-ALP,

which can be easily detected by amperometry. An incon-

venience of this system, with two different biotinylated

oligonucleotides in the sandwich system, is the requirement

of a blocking step with biotin in between, in order to avoid

the adsorption of the second biotinylated strand on the

streptavidin surface. The microfluidic cartridge is commer-

cialized by DiagnoSwiss SA under the ImmuChipTM name.

The chip consists of eight microchannels etched in poly-

amide. Each microchannel contains eight carbon screen-

printed working electrodes and in the upper part of the

channel screen-printed counter and reference electrodes.

The amperometric detection limit obtained from the ALP

label was 0.2 nM, obtaining a negligible response from the

non-complementary sequences.

A similar sensoric platform was used in the chip

developed by Gabig-Ciminska et al. [12], who also use a

sandwich format labeled with ALP on streptavidin magnetic

beads. This device was developed for the analysis of 16S

rDNA in Escherichia coli extracts. The chip consists of four

pairs of gold IDE patterned with lift-off on silicon wafer and

encapsulated with silicon oxynitride. The target was ampli-

fied by PCR and detected amperometrically on the chip. The

selectivity of the capture probe and the non-specific

adsorption of the label were tested, obtaining a good anti-

fouling with BSA. The detection limit reported with this

chip was 1 nM.

3 Voltammetric DNA biosensors

Voltammetric technique determines the amount of matter

transformed during a redox reaction when the potential

varies with time in a pre-determined manner and the

current is measured as a function of potential. Redox

molecules are detected with this technique [13].Figure 1. Schematic of the amperometric detection mechanismbased on a sandwich conformation.

Electrophoresis 2009, 30, 3386–33973388 M. Mir et al.


3.1 Voltammetric sandwich

The e-SENSORTM reported from Motorola in 2001 [14] is a

voltammetric sandwich-based sensor that six years later was

incorporated in a fully integrated biochip [15]. This is a full-

integrated cartridge that comprises different units for pre-

treatment and detection of the sample. This LOC integrates

a mixing unit for rare cell capture using immuno-magnetic

separation, which is a useful tool in order to enrich rare

target cells for a large volume sample, a single chamber unit

for cell pre-concentration, purification, lyses and PCR, in

order to simplify the chip, and a DNA microarray detection

chamber. The chip was fabricated in plastic, with a

patterned printed circuit board and a 4� 4 gold microarray

chip. It was tested with real whole blood sample for the

detection of hematochromatosis-associated SNP. All the

sequence of processes, from the blood injection to the signal

readout, takes about 3 h (90 min for the PCR and 60 min for

the detection). The hybridization was detected though a

sandwich assay, using a second probe ferrocene labeled,

detected by alternating current voltammetry.

Magnetic streptavidin-coated beads are used to concen-

trate and detect DNA in a sandwich format also in other

studies [16]. In the work of Baeumner et al. a redox molecule

entrapped in a liposome capsule serves as a label to the

second probe, which recognizes specifically the synthetic

DNA sequence of Dengue virus serotype 3. The hybridized

target flows through soft-lithographed PDMS microchannel

attracted by a magnet to an interdigitated array patterned by

lift-off on glass substrate. Once the magnet immobilizes the

target, the lysis of the liposome is induced with b-octylglu-

copyranoside, realizing the redox molecule entrapped that is

voltammperometrically detected. The encapsulation capacity

of the liposomes brings the advantage of having a high

amount of redox molecules that can be detected in each

hybridized duplex. Oppositely only one redox molecule perhybridized target can be detected in other systems. Never-

theless, this sandwich with liposome label reports a similar

limit of detection (10 nM) as the system with a single label

per duplex. Moreover, the structure of the liposome

capsule could be unstable in certain environments, so in a

complex real sample matrix this kind of labels could not be

efficient.

3.2 Redox intercalators

In order to remove the needs of a label for the hybridization

detection, which is introduced mainly though a sandwich

format or by PCR, redox DNA intercalator molecules can be

used. These kinds of redox molecules have the appropriate

size and chemical nature (polycyclic, aromatic and planar)

which makes them fit in between base pairs of DNA. Thus,

once the double strand is formed, after the target

hybridization with the capture probe, the intercalators are

entrapped in the double helix. These redox molecules are

detected by voltammetric techniques.

Simultaneous DNA amplification by PCR and subse-

quent electrochemical detection was performed in a micro-

well [17]. Platinum temperature sensors, heaters, and gold

electrodes were patterned on top of the silicon substrate that

was covered with a bottom glass substrate for sealing the

chamber. In a 8 mL silicon chamber an asymmetric ampli-

fication was carried out that brings thousands of single

stranded target sequence, which are incubated for 1 h in the

same chamber in order to facilitate its hybridization with

complementary capture probes immobilized on gold elec-

trodes patterned on the chamber top. In order to get rid of

any unhybridized molecule, the chamber was rinsed and

washed. The hybridization was detected with differential

pulse voltammetry by means of redox molecules intercalated

into the double helix. The double function of the chamber

facilitates and reduces cost of the device fabrication and

minimizes the size of the chip.

The hybridization of the target with the capture

probe can also be detected in solution, without previous

immobilization of these molecules [18]. The capture

probe and the synthetic DNA target are introduced

separately into a microchannel, where both flow and

hybridize. The created double strand of DNA was inter-

calated with a redox molecule and it was detected by cyclic

voltamperometry when the double strand flows on the area

where the electrodes were patterned. The hybridization

selectivity was tested with a mismatched sequence obtaining

20% of the signal. Although this chip removes the problem

associated with the immobilization of the DNA receptor,

this configuration does not allow a washing step after

hybridization, which is very important in real sample

with complex matrix that complicates the hybridization

detection.

3.3 Electronic beacon

The e-DNA sensor reported by Fan et al., in 2003 [19] was

applied in microfluidic devices five years later [20]. The

ferrocene-labeled DNA stem-loop structure used in the

e-DNA sensor changes its conformation due to target

hybridization, which alters the electron transfer tunneling

distance between the electrode and the ferrocene label that

can be detected by differential pulse voltammetry (Fig. 2). In

this chip three different ferrocene-thiolated beacon probes

were attached on the gold array by means of electronic site-

selective deposition. After use, the electrodes were almost

totally (97%) regenerated by positive potential sweep

desorption. The gold array was patterned on glass and the

microfluidic channels were incorporated in a PDMS slide.

The electrochemical beacon configuration permits a

reagentless and label-free system. Furthermore, the non-

complementary neighboring capture probes showed little

effect on the signal (3%), demonstrating that it is a specific

sensor. However, a background signal of about 60% from

the ferrocene signal remains after hybridization of synthetic

oligonucleotide target.



3.4 Guanine base detection

The direct electrochemical analysis of the oligonucleotide

strand can bealso performed, since guanine bases are

electroactive molecules.

A separation and pre-concentration process integrated

in an electrochemical DNA sensor based on this principle

was developed by Shiddiky and Shim [21]. In order to

improve the separation and detection of DNA molecules in

the Teflon microchannels, colloidal gold (CG) nanoparticles

were used in both processes. Hydroxypropyl cellulose mixed

with CG nanoparticles creates a matrix into the micro-

channel and with voltage application separates the different

sizes of DNA and concentrates the sample. In the detection

step CG nanoparticles entrapped in a conductive polymer

helps in the current conduction for the DNA strand to the

electrode surface, which allows the direct label-free detection

of the hybridization. The CG current amplification and the

target concentration permit low detection limits of 700 pM.

Carbon nanotubes (CNT) are nanowires (NW) of high

conductive surface that provides a flexible surface chemistry.

In the majority of CNT-based sensors reported, the CNT

forms a three-dimensional structure of long carbon wires,

bringing a large graphite surface to functionalize but also a

high capacitive discharging current/surface area. In the

article reported by Li et al. [22] the CNT are encapsulated

with SiO2, leaving only the tip of the tubes exposed to the

solution. This procedure forms a nanodisc electrode surface,

which was repeated for each spots on an array and inte-

grated in a chip device. The amino-capture probe is inter-

acted with the carboxylic groups on the CNT surface and

hybridized with label-free PCR-amplified DNA target. The

hybridization is detected by voltamperometry through the

direct oxidation of guanine bases. These bases on the

capture probe are replaced with non-electroactive inosines to

eliminate the redox background from the probes. Ru(bpy)321

mediator is introduced in order to amplify guanine oxida-

tion based on an electrocatalytic mechanism with the

guanine bases that amplify the oxidation detection (Fig. 3).

The sensor was tested with non-complementary strands,

which brings lower response.

Another example of direct oxidation detection of DNA

strand was reported in ref. [23]. The sensor was mounted in

a glass substrate bounded with PDMS microchannels

network. Cyclic voltammetry was used in this case in order

to measure the oxidation of guanine bases on the hybridized

target after a long hybridization incubation of 2 h 30 min

with a thiolated capture probe immobilized on gold elec-

trodes. A mixture of Ru(NH3)631/Fe(CN)6

3 is loaded into the

channel in order to catalyze the guanine oxidation. The

hybridization even at 500 nM of target is revealed with a

weak increase on the cyclic voltamperometry ruthenium

peak. The signal obtained from the ruthenium mediator

after and before hybridization is very similar, so a method

that amplified this differences as well as an optimization of

hybridization time is necessary.

3.5 PCR redox labeling

A definite oligonucleotide sequence is specifically amplified

using the appropriate primers with PCR. Therefore, DNA

not containing this specific sequence, which could be

determinant for the diagnosis of a disease, is not amplified.

Based on that a microfabricated electrocapillary electrophor-

esis chip that contains an electrochemical detector was

carried out for the SNP typing assay of C282Y substitution

diagnostic for hereditary hemochromatosis [24]. The PCR

assay was performed with a ferrocene-labeled primer, so if

the target is specific to these primers, the target will be

Figure 2. Schematic representation of an electrochemicalbeacon DNA sensor.

Figure 3. Mechanism of electrochemical amplification detectionof guanine bases with Ru(bpy)3

21 mediator.



amplified and labeled with ferrocene. The remainder of

reagents from the PCR are separated by CE on polyacryla-

mide-coated channels using a hydroxyethylcellulose sieving

matrix at a potential of 400 V/cm. In order to minimize the

effect of the electrophoresis electric field on electrochemical

detection, the working electrodes were placed between

200 mm from the separation channel exit, where the

ferrocene label was detected with voltamperometric techni-

ques. The absence of response obtained with a mutated

sequence corroborates the selectivity of the sensor.

Based on the same electrophoretic separation and

detection principle, a PDMS microchip was developed for

159 bp PCR amplicon detection [25]. In this case the ferro-

cene label was incorporated after the PCR amplification. The

label was attached to a single base that is hybridized to the

amplified double helix. This system is more time consum-

ing than the previous one because this needs an extra step

for the hybridization of a single base labeled that is not

100% efficient. This step could reduce the number of targets

labeled, compared with the labeling through the PCR

primer, with a subsequent decrease in sensitivity.

4 Potentiometric DNA biosensors

Potentiometric biosensors are based on charge potential

changes at the sensing electrode with respect to the

reference one that is usually translated in a voltage change.

On these measurements current should be kept nearly zero.

DNA potentiometric detectors based on current

measurements that result in a charge accumulation effect at

the gate electrode of transistor devices by specific binding of

DNA molecules are also referred as field-effect DNA

biosensors. These DNA biosensors are particularly attractive

since they can be fabricated in very small sizes and

massively integrated in LOC devices in silicon manufactur-

ing technologies that could integrate many other electronic

elements.

Pure electrochemical field-effect biosensors are mainly

metal-oxide-semiconductor structures and can be classified

into electrolyte–insulator–semiconductor and ion-selective

field-effect transistors (FET) devices Fig. 4-A. A particular

type of those are NW detectors including CNTs. NW are

long tubes with diameters of a few nanometers and due to

their small dimensions their conductivity behavior is

strongly dependent on external charges and biocompounds.

Surface charge of the NW will strongly influence the current

circulating through it.

4.1 Field-effect DNA biosensors

Sakata et al. developed a genetic FET chip based on the

potentiometric detection of hybridization and intercalation

on its gate [26]. Four FET and a temperature sensor were

integrated in a chip area of 5 mm by 5 mm. The detection of

charge density change as a result of hybridization using

genetic FET was demonstrated. Oligonucleotide probes were

immobilized in a Si3N4 gate surface and the electrical output

signal was obtained directly without any labeling. The

threshold voltage of the transistor shifted along the X-axis by

78 mV in the positive direction. This positive shift was due

to negative charges of the oligonucleotide probes and to

various intrinsic charges induced on the gate surface during

Figure 4. Representation of several field-effect DNA biosensors. (A) The SiO2, alternatively Si3N4, layer passivating the surface betweenthe electrodes on the gate of the ion-selective FET sensor can be functionalized with complementary target oligonucleotides in order toachieve specific recognition. (B) A single CNT acts as a conduction channel between two electrodes in a DNA NTFET sensor. The CNTcan be functionalized with immobilized oligonucleotides. (C) Alternatively a CNT network with immobilized synthetic oligonucleotidescould be used instead of a single NT.



immobilization process steps. When the complementary

target oligonucleotides were introduced on the gate surface,

the threshold voltage shifted in the positive direction 11 mV

further. This was due to the negative charges of comple-

mentary oligonucleotide as a result of hybridization at the

gate surface. Nearly all the oligonucleotide probes were

considered to be hybridized with the target DNA since a

comparatively high concentration of target DNA (100 mM)

was used and hybridization time was sufficient [26].

In 2006, Barbaro et al. fabricated tested an integrated

field-effect device for fully electronic DNA detection [27].

The sensors part of the chip was realized in a standard

0.8 mm CMOS process by Austria Microsystems and

consisted of 16 sensors divided in two clusters to test

detection capabilities. The sensing surfaces of both clusters

were activated with two different oligonucleotide probes.

The microfluidics structure was build on top of the sensors

with PDMS and the sensors were previously directly bonded

to a printed circuit board having overall area dimensions of a

microscope cover slip. Hybridization tests were performed

using two different solutions containing the targets

dissolved in a low ionic strength buffer in order to minimize

electrical noise. The two solutions were introduced using

syringes in two small reaction chambers of the microfluidic

structure that contained the clusters of sensors. The

temperature of the chip was maintained at 501C and the

electrical signals coming from the biosensors were analyzed

by a computer. A differential threshold voltage was calcu-

lated using one sensor in each cluster. The computed

threshold voltage of each sensor in the cluster used as a

reference was subtracted from the one obtained in the

corresponding sensor in the active cluster. This methodol-

ogy allowed the elimination of changes due to a global

cause, such as rinsing procedure, that were canceled by

differentiation since they affected all sensors at the same

time. Specific single sensor changes, such as successful

hybridization, were detected as a robust differential signal.

Completion of hybridization reaction was detected on eight

pairs of sensors for the first target and for five out-of eight

for the second. It was believed that biological reaction

statistical fluctuations affected the result.

The device sensors were designed to be fully compatible

with the standard commercial CMOS processes demon-

strating the feasibility of realizing low-cost, large-scale

integrated, and fast, DNA electronic detectors.

4.2 CNT and nanowire field-effect DNA biosensors

In the nanotube field-effect transistor (NTFET) a single

wall nanotube (SWNT), Fig. 4(B), or a network of those,

Fig. 4(C), acts as a conduction channel between two

electrodes. The current that flows through the channel is

modulated by an electric field applied at the gate electrode.

The NTFET working principles are still under debate [28].

Variations on the electrode to CNT contact resistance

rather than on the channel conductance are reported as the

response mechanism by several works. The CNT transistors

would work as unconventional Schottky barrier transistors

[29] and their performance will depend on the source–drain

electrodes geometry and conductive material properties as

well as the nanotube diameter. CNT with diameters over

1.4 nm and electrodes with the lowest metal work function

have been reported as having the best NTFET performances

[30]. A shorter tunneling barrier will produce a device with

higher maximum ON-state current and in a sensing system

an electron donating molecule could be able to further lower

the electrode work function [28]. However, in liquid-gated

NTFETs the nanotubes are very sensitive to the electrostatic

environment and small disturbances caused by biomole-

cules can lead to significant changes of the device conduc-

tance [31]. Furthermore, the device capacitance changes will

also slightly influence the NTFET characteristics. In this

case, local permittivity decreases in relation to the electrolyte

solution due to molecular adsorption. Most of the results

reported in literature when using these devices as biomo-

lecular sensors in liquid environments seem to be related to

the tunneling barrier and the electrostatic gating. The

passivation of the metal–NT contact would allow more

controllable biosensors since the Schottky barrier effect

seems to be less reproducible and consistent.

However, NTFET presents enhanced electrical proper-

ties in respect to other field-effect classical devices and it has

several structural characteristics that make them very

suitable as sensing devices, such as having a large surface

packed of atoms and being easy to functionalize [28–34].

NTFET devices sensitivity and surface-selective chemistry

surface show them to be very good candidates for DNA

electrochemical sensing. Furthermore, recent studies have

shown that SWNT adsorb DNA molecules that probably

produce a change in its conformation and its electronic

structure [35, 36].

A CNT network with immobilized synthetic oligonu-

cleotides NTFET chip working as a selective detector of DNA

immobilization and hybridization was reported by Star et al.[37]. Interdigitated NTFET devices were fabricated by using

chemical vapor deposition at 9001C to grow SWNT on top of

doped Si wafers with SiO2 at their surface. Dispersed iron

nanoparticles were used as growth promoter. Evaporated

Ti–Au films were patterned on top of the nanotubes by

standard photolithography to create the electrical leads.

H63D SNP discrimination, in a gene responsible for

hereditary hemochromatosis, was accomplished by moni-

toring the NTFET electronic transfer characteristics chan-

ges. SNP discrimination was done in the presence of 5 mg/

mL non-homologous DNA, similar to the concentration of

DNA in 1 mL of blood. The use of covalently attached DNA

capture probes, surfactants to lower background noise, and

microfluidic on-chip structures to improve the sample-

delivery method were proposed as further improvements.

Tang et al. [38] also reported the fabrication and test of

NTFET DNA sensors and studied their sensing mechan-

isms. In this case, the patterned electrodes in each sensing

chip were made of Au–Ti and covered with a self-assembled



monolayer of mercaptohexanol. They found that DNA

hybridization on the electrodes was mainly responsible for

the conductance changes due to modulation of the electrical

alignment between the SWNT and the gold contacts. It

induced a change in the metal work function of the device

electrodes. The SWNT translated the hybridization on top of

Au electrodes into an electrical signal. Recent results [39]

also suggested that the device electrode may play a signifi-

cant role in the observed NTFET response.

The proposed fabrication method was readily scalable to

high-density sensor arrays to be integrated in LOC devices

incorporating microfuidics to prepare and deliver the samples.

Alternatively and similarly to the nanotubes approach

silicon nanowires (SiNW) can be also used in FET DNA

biochemical sensing devices. These nanostructures can be

tuned in a quite reproducible manner by controlling their

diameter, dopand concentrations, and surface modification,

due to the insulating native oxide around the NW [40]. A

recent article [41] has shown a delicate trade-off between the

sensitivity of the SiNW sensor and the stability of the recep-

tor–target molecule as a function of ion concentration. The

sensitivity of these nanosensors will increase when doping,

diameter and length of the NW are reduced. The amount of

induced charge in these NW is set by the dielectric properties

of the analyzing media due to electrostatics effect.

SiNW FET sensors functionalized with peptide nucleic

acid receptors were shown to distinguish between mutated and

wild type cystic fibrosis transmembrane receptor gene by NW

conductance changes [42]. An array of SiNW FET sensors for

multiplexed detection of cancer markers was presented in 2005

showing very high sensitivity and selectivity [43]. It successfully

detected low concentrations (0.9 pg/mL) in undiluted serum

samples, having the cancer marker surpassed by proteins

100 billion times. Furthermore, in the same work, a similar

device was able to measure telomerase binding and activity

that have been related to many cancer processes, without

amplification in samples having only ten cells. Recently,

detection of DNA methylation without PCR amplification and

bisulfite treatment using NW has been demonstrated [44]. The

linear wires, with diameters ranging from 28 to 100 nm, were

grown at precise contact points using existing semiconductor

manufacturing techniques. Methylated promoter was recog-

nized by monoclonal anti-5-methylcytosine antibodies immo-

bilized on the NW surface. The best results were obtained with

the 28 nm NWs achieving significant detection sensitivity at

2.5� 10�19 mol DNA. No false positives were detected in non-

methylated DNA samples up to 2.5� 10�15 mol.

It is expectable that many applications of these novel

sensors to the LOC technology will appear in the fore

coming years.

5 Impedimetric DNA biosensors

Impedimetric transduction is based on sensing the electrical

resistance and reactance of the evaluated particular medium

delimitated by the measuring electrodes.

Among the different methods used in this type of

sensors, electrochemical impedance spectroscopy is a very

well-known technique that dates back to the mid seventies.

On the other hand, the advent of nanotechnology and its

application to sensor and LOC technology have brought

more recently a particular, relatively new and very promis-

ing, type of impedance DNA biosensors: the nanopores.

5.1 Impedance spectroscopy detection

Electrochemical impedance spectroscopy is measured by

applying a small excitation signal of alternating current

potential to an electrochemical cell and measuring the

circuit ability to resist the flow of electrical current through

the cell. This method is highly sensitive and is also able to

detect changes on the electrode surface upon biomolecule

interactions [45].

A microfluidic device that incorporates CE for biomo-

lecule separation used this technique to detect 8-hydroxy-

deoxyguanosine DNA adduct, which is a frequently used

biomarker reported on the oxidative stress [46]. Both elec-

trodes in the electrophoresis and in the detection chamber

were electroplated with palladium in order to improve the

separation efficiency and the detection limit. However, the

concentration limit achieved was about 100 nM, which is

high in comparison with most of the electrochemical chip

devices summarized in this review.

It is important to notice that changes on the buffer or on

the distance between the working and the reference elec-

trode will bring changes on the measurement as a result of

the high sensitivity of the impedance technique. A glass

microchip with integrated microelectrodes was fabricated

for the purpose of studying these effects in a work done by

Park et al. [47]. The chip was tested with synthetic lambda

DNA as a sample biomolecule.

The majority of impedimetric sensors, as these

previously reported, detect Faradaic impedance of the bior-

eceptor immobilized on the electrode surface. Examples of

non-Faradaic impedance detection of biomolecule interac-

tions are less reported in literature and mainly consist in

attaching the capture probe in between the two electrodes.

In this kind of sensors, IDE configurations are commonly

used in order to have a narrow gap between the electrodes,

since the distance between electrodes is directly related to

the sensitivity of the sensor. The electric field generated in

one of the IDEs is distributed in the area where the target

interacts with the bioreceptor, creating a change of the

generated electric field, which is detected on the second

IDE. The ionic media is a drawback in this kind of sensors,

because of the pre-dominant electrical spreading resistance

of the solution. The group of Gris [48] overcomes this

problem optimizing the washing step after the target

hybridization in order to remove all the ions from the

solution. Their chip was composed of four IDE fabricated

with lift-off photolithography integrated in a PDMS

patterned with microfluidic channels and mechanically



clamped in between two slides of PMMA. The four sensing

areas were functionalized externally through silane chem-

istry on glass with different capture probes in order to check

the non-specific adsorption of the sample in the absence of

capture probe and with mismatched sequences. The chip

was first tested with synthetic DNA target obtaining a

detection limit of 10 nM. Afterwards the capability of the

chip with real DNA samples was demonstrated with the

detection of PCR-amplified DNA from S. choleraesuis.

5.2 Nanopore DNA detection

Nanopores, or holes of dimensions of a few nanometers,

have shown a huge potential to become the next generation

of DNA detectors and sequencers. A recent review about the

potential and the challenges of this DNA impedimetric

detection technique has been recently published [49]. These

novel DNA detectors are based in a very well known and

highly used technique to count and characterize particles.

This technique is named the Coulter counter (Beckman

Coulter, Fullerton, CA, USA) and it is based in electrically

driving a particle through a fluidic restriction (Fig. 5), the

pore, keeping track of the current changes at given applied

constant voltage [50]. This and other similar techniques are

broadly known as resistive-pulse sensing, a detailed general

review has also been published [51]. Since the inner

diameter of nanopores is on the same scale as many single

molecules it can be used as a restriction to monitor them.

The ionic current through the nanopore of a suitable

diameter is modulated by the nucleotides of RNA, or DNA,

strands as they are electrophoretically driven through it at a

constant determined voltage [52].

Construction of nanopore structures was originally

achieved by using naturally occurring protein pores in a

lipid bilayer [53, 54]. Interest raised and research done

during the last ten years have allowed to achieve them in

more stable solid-state materials [55, 56], graphene [57], or

polymers, such as polycarbonate [58] or Kapton [59].

However, the most studied and commonly used nano-

pore for sensing purposes is still a-hemolysin (aHL) protein

that spontaneously inserts itself into lipid membranes,

Fig. 5(A). This toxin is still functional at high temperatures

and spontaneously creates transmembrane channels of

1.4 nm diameter at its narrowest point. Owing to its

dimension, it allows the passage of ssDNA that unravels

locally forming an ordered nucleotide cue while crossing the

pore [54].

As a drawback of the technique, it was soon realized that

ion–current blockades at the aHL nanopore was a conse-

quence of the effect of several nucleotides [60] since the

RNA bases translocation at the narrowest space occurs too

fast [61]. Since then, different alternatives have been used to

modify this protein to use it as a nanopore DNA sensor. By

chemically engineering these proteins it is possible to probe

complimentary DNA hybridization. Nucleotide oligomer

covalent attachment to the lumen of aHL binds to the

matching sequence delaying the pore translocation and

achieving a sequence-specific detection sensor of single-base

resolution [62]. Mutated aHL pore having a higher positive

charge have been created in order to improve the frequency

of ssDNA capture from solution and enhance the sensor

sensitivity [63]. A recent article showed that the nanopore

with a covalently attached adapter is able to identify

nucleoside 50-monophosphate molecules with accuracies

averaging 99.8% [64]. Furthermore, methylated cytosine can

also be distinguished from the standard DNA bases. The

Oxford Nanopore Technologies is currently integrating this

sensing method in a complete system to analyze epigenic

modifications.

Moreover, studies have shown that unmodified aHL

interacts with DNA hairpin molecules and the computa-

tional analysis of the complex signature allows the discri-

mination among DNA hairpins at single nucleotide

resolution [65].

The use of protein-based nanopores has several limita-

tions in front of its man-made counterparts. They have

limited chemical and thermal stability, relatively fixed size

and shape, they are not easy to integrate on microdevices

and to modify to change their DNA translocation and affi-

nity properties. Furthermore, 5.4 kb lengths of ssDNA have

been threaded through solid-state nanopores in front of the

25 kb achieved by the protein engineered ones [49].

Artificial nanopores, Fig. 5(B), could be massively inte-

grated in a microfluidics chip and the idea of integrating a

nanopore sensor system in a Micro Total Analysis systems

device could make complete DNA sequencing cheap, highly

Figure 5. Two basic types of nanopores are currently used: (A)aHL, from the bacterium Staphylococcus aureus, and mutationsof it, is used as a natural occurring nanopore and it is limitatedby its fixed dimensions and shape. (B) Artificial nanopores madeof semiconductor industry, graphene, or polymeric, materialscan be tuned to achieve desired properties, and are more stableand easy to integrate in LOC devices.



https://www.researchgate.net/publication/37449397_Particle_Size_Measurement?el=1_x_8&enrichId=rgreq-428e052acdabbdbc06bcb0d9aeb5e1d5-XXX&enrichSource=Y292ZXJQYWdlOzI2ODY5MDcyO0FTOjEwNjYzMjcwNDk1NDM2OEAxNDAyNDM0NjI2NjY0

parallelizable, portable and fast. Regular microfabrication

materials, as silicon oxide and silicon nitride, have been

proven suitable to build useful DNA sequence nanopores

using clean-room techniques as ion beam sculpting or

e-beam drilling [55, 56].

6 Concluding remarks and futureprospects

This review summarizes the strategies carried out by

different research groups, in order to integrate electroche-

mical DNA biosensors in microfluidic cartridges and some

of these include various sample pre-treatments, with the

purpose of miniaturize all the steps necessary in a DNA

analysis laboratory. Table 1 summarizes the main articles

described in this review, in order to have an easier overview

on these. Although, some techniques already have shown a

great deal of maturity, some other very promising ones are

still on their youth and many things have to be further

improved and explored in order to exploit the best of their

capabilities. This is especially true for the nanotechnology

advances related techniques such as field-effect NW, and

particularly CNT, and impedimetric nanopore-based

approaches.

NW and CNT DNA biosensors working principles are

still not fully understood and this brings many doubts about

results repeatability and stability. A great effort to under-

stand the basic principles behind them and to develop a

theory strongly individuating and isolating the different

affecting physical phenomena in their performance is

needed before being able to use them on commercial devi-

ces.

On the other hand, sequencing of single bases by means

of nanopore molecular sensors has proven to be out of reach

for the moment since the oligonucleotide strands tend to

cross the nanopore at very high velocities. Combination of

these techniques with other oligonucleotide manipulating

methods and tuning of the nanopore properties by applying

biochemical engineering and functionalization of the pore,

or by using other fabrication materials and techniques,

could bring a solution to this problem in the next years.

Furthermore, fabrication of these structures integrated on

LOC devices is still not straightforward bringing an addi-

tional challenge to be overcome by technology in order to

massively use this type of sensors.

Finally, the great advantage of integrating DNA biosen-

sors into an LOC device should be the reduction in sample

and reagents volumes. However, to fully achieve this and

bring these devices into profitable commercialization it is

needed to integrate the sample preparation protocol into the

LOC device. This introduces many other challenges such as

the a priori smart design, simplification, and combination of

the different protocol steps, or the interconnection of differ-

Table 1. Summary of integrated electrochemical DNA biosensors for LOC devices. LOC analysis time considers the time of analysis

inside the LOC and total analysis time considers the LOC analysis time adding the pre-treatment time required outside the LOC

Detection technique

and references

On chip sample treatment LOC analysis

time (min)

Total analysis

time (min)

Reprodu-

cibility

Sensitivity Application

Amperometry [6] None 190 190 NR 10 nM Synthetic DNA

Amperometry [9, 10] External purification 45 105 75% 12 pM Bacterial 16S rRNA

Amperometry [11] None 55 55 88% 200 pM DNA specific to hazelnut allergent

Amperometry [12] None 240 240 86% 1 nM 16S rDNA E.coli

Voltammetry [15] Concentration, purification,

lysis and PCR

150 150 NR NR Hematochromatosis SNPs

Voltammetry [16] None 20 20 NR 10 nM DNA specific to Dengue virus

serotype

Voltammetry [17] PCR 130 130 NR NR 5S rRNA Fritillaria thunbergii

Voltammetry [20] None 25 25 98% 400 nM Human H1N1 and avian H5N1

Voltammetry [21] Concentration & separation 10 10 NR 100 fM Synthetic DNA

Voltammetry [23] None 150 150 NR 500 nM Synthetic DNA

Voltammetry [24] External PCR 10 85 NR NR Hematochromatosis SNPs

Voltammetry [25] External PCR 200 276 NR NR SNPs

Potentiometry [26] None 720 720 NR 100 mM Genetic analysis

Potentiometry [37] None 60 60 NR 100 pM Hemochromatosis SNPs

Potentiometry [38] None 140 140 NR 100 nM Synthetic DNA

Potentiometry [42] None 3.3 3.3 NR 10 fM Genetic risk assessment by

mutation screening

Potentiometry [43] None 3.3 3.3 NR 24 fM Detection of cancer markers

Impedance [46] Electrophoresis separation 1.7 1.7 NR 100 nM Urinary 8-hydroxy-deoxyguanosin

biomarker detection

Impedance [48] None 60 60 NR 10 nM Detection of pathogen DNA

NR, not reported.



ent blocks and its dead fluid volumes reduction. Further-

more, the choice of the right LOC material for the application

as well as the reduction in the final production costs are also

issues that sometimes escape the mind of scientists and that

can be great impediments to the onset of industrial and social

interest. However, these limitations are the same that are

found in other transduction methods, such as the optical,

once they are evaluated to be integrated on an LOC device and

they are being resolved in most of the cases for each parti-

cular application. Moreover, electrochemical genetic sensors

present the advantage of a complete electronic read-out that

reduces the cost of external interfaces and facilitates the

choice of the LOC device material that would not have the

need of highly demanding optical properties.

It is clear by the end of this review that electrochemical

oligonucleotide detection presents several advantages over

other methods suitable to be integrated on LOC devices.

Several commercial applications are already on the market

and there is a nice prospective for the appearance of new

products based on these technologies in the fore coming

years despite the actual limitations.

The authors have declared no conflict of interest.

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Integrated electrochemical DNA biosensors for lab-on-a-chip devices

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