A Fully Electronic Label-Free DNA Sensor Chip

IEEE SENSORS JOURNAL, VOL. 7, NO. 4, APRIL 2007 577

A Fully Electronic Label-Free DNA Sensor ChipClaudio Stagni, Carlotta Guiducci, Luca Benini, Fellow, IEEE, Bruno Riccò, Fellow, IEEE, Sandro Carrara,

Christian Paulus, Meinrad Schienle, and Roland Thewes

Abstract—This paper presents a microfabricated DNA chip forfully electronic, label-free DNA recognition based on capacitancemeasurements. The chip has been fabricated in 0.5- m CMOStechnology and it features an array of individually addressablesensing sites consisting of pairs of gold electrodes and addressinglogic. Read-out circuitry is built externally using standard compo-nents to provide increased experimental flexibility. The chip hasbeen electrically characterized and tested with various solutionscontaining DNA samples. Significant capacitance variations due toDNA hybridization have been measured, thus showing that the ap-proach represents a viable solution for a single chip DNA sensorarray.

Index Terms—Capacitance measurement, DNA sensor arrays,hybridization label-free detection.

I. INTRODUCTION

I N THE LAST decade, miniaturized arrays for gene-basedtests (known as DNA microarrays) have been introduced in

genetic research. These are fabricated on glass or quartz slideswhere different DNA probe molecules are immobilized in atwo-dimensional array of small sites. Some of these devices,implemented with photolithographic techniques, are able to testa whole genome, with densities of a million sites per squarecentimeter [1]. The DNA strands of the organism under testare preliminary marked with fluorescent molecules. Their pres-ence at specific sites of the array is measured by an opticalscanner or a fluorescence microscope, after binding to matchingprobe molecules. These devices have also been employed forpopulation genotyping [2] and research on cancer predisposi-tion [3]. Nevertheless, the high cost of the scanner and the pro-cessing steps required to tag the samples with markers (alsocalled labels), pose critical limits to the use of these tools inpoint-of-care analysis. For these reasons, a large effort is de-voted to the development of devices suitable for low cost massproduction and ease to be used not only in highly specializedlaboratories. A single-chip solution, implementing a detection

Manuscript received March 6, 2006; revised October 11, 2006. The associateeditor coordinating the review of this paper and approving it for publication wasDr. Alvin Crumbliss.

C. Stagni, L. Benini, and B. Riccò are with the Department of ElectricalEngineering and Computer Science (DEIS), University of Bologna, 40136Bologna, Italy (e-mail: [email protected]; [email protected];[email protected]; [email protected]).

C. Guiducci is with the Laboratoire de Nanobiophysique-ESPCI 10, 75005Paris, France (e-mail: [email protected]).

S. Carrara is with the Dipartimento di Biochimica, 40126 Bologna, Italy(e-mail: [email protected]).

C. Paulus and M. Schienle are with Siemens AG, 81739 Munich, Germany(e-mail: [email protected]; [email protected]).

R. Thewes is with Qimonda AG, 85579 Neubiberg, Germany (e-mail: [email protected]).

Digital Object Identifier 10.1109/JSEN.2007.891990

technique with direct electrical read-out and avoiding labelingof the DNA target molecules, would significantly enhance porta-bility, as well as high parallelism, on-site sensing, and data pro-cessing.

This paper provides a contribution in this direction since itpresents a CMOS DNA-chip based on a label-free fully elec-tronic capacitance measurement technique. The chip features anumber of sensing sites, each containing two gold electrodes[4] that can be independently selected by means of on-chip ad-dressing circuitry, while the read-out electronics has been real-ized essentially by means of external standard components inorder to exploit the advantages of large experimental flexibility.The surface of the electrodes is bio-modified (functionalized)by covalent binding of single-stranded DNA probes. Duringin-field operation, the capture of complementary DNA strandsfrom the liquid samples is signaled by a variation of the capac-itance between electrode pairs, which can be measured by con-necting them (through on-chip addressing logic) to an externalcapacitance-measurement circuit.

The chip is bonded on a printed circuit board (PCB) and isconnected to external processing and storage devices, resultingin a complete analysis system that has been characterized andtested for DNA recognition. To this end, the capacitor elec-trodes have been functionalized and their capacitance variationhas been measured after exposition to both complementary andnonmatching target molecules (to test specificity). Reversibilityhas also been tested by measuring sensor capacitance before andafter sensing molecule regeneration, obtained by temperaturetreatment. The results demonstrate that a DNA sensor based onthe electronic detection technique described in this paper canbe fabricated as a single chip with standard CMOS technologyaugmented with gold deposition for the capacitor electrodes.

This paper is organized as follows. In Section II, anoverview about state-of-the-art electronic biosensors is given.In Section III, the capacitive detection approach used hereis introduced. Section IV describes the chip architecture andoperation. In Section V, the measurement setup is illustrated.In Section VI, experimental results are presented. Finally,conclusions are drawn in Section VII.

II. RELATED WORK

DNA microarrays allow highly parallel and low cost analysis.In fact, they exploit the capability to fabricate a large number ofminiaturized detection sites on a substrate and to extract infor-mation after exposition to the target DNA to be detected/rec-ognized. Each site is specifically bio-functionalized by meansof DNA probe molecules of known sequence immobilized onits surface. Target molecules in the sample solution bind onlyto probes with complementary sequences (hybridization): thus,their presence at specific sites reveals their composition.

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578 IEEE SENSORS JOURNAL, VOL. 7, NO. 4, APRIL 2007

Most of current microarrays require optically active labels at-tached to the target DNA and fluorescence sensing. Some in-novative microarrays employ electrochemical labels (redox-ac-tive molecules or enzymes) that produce an electrical currentthrough conductive electrodes in case of hybridization [5]–[7].

Label-free techniques are intensively investigated, sincethey avoid expensive reagents and pre-treatment steps. To thisend, recently, several approaches have been proposed, basedon piezoelectric materials [8], [9], microfabricated cantilevers[10], [11]. Other techniques detect the variations of electricalproperties of electrode–solution interfaces induced by DNArecognition. Within this category, a distinction can be madebetween devices where sites exploit passive components orsemiconductor sensors [12]–[17].

The approach of interest in this work is based on interface ca-pacitance measurements, whose conventional implementationexploits a three-electrodes setup allowing biofunctionalizedgold electrodes to be investigated with impedance spectroscopy[18], [19].

Much simpler two-electrodes configurations, such as thatused in this work, have also attracted significant interest. Forexample, a two-electrode setup using 2.4-mm gold and aAg–AgCl counter-electrode has been found to work properlyin capacitive detection at 20 Hz [20]. As far as integrablecapacitance methods are concerned, it has already been shownthat DNA hybridization can be detected by measuring interfacecapacitance with a system making use of two (gold) elec-trodes only (i.e., without the use of a reference terminal) [21].Furthermore, more recently it has been shown that reliablemeasurements can be done exploiting passive microfabricatedelectrodes [22]. Arrays made of passive electrodes are pad-lim-ited, as each electrode pair requires two pads for connecting toexternal measurement circuits.

In such a context, this work represents an advance in thestate of the art in that it presents a silicon chip featuring mi-crometric capacitors and on-chip logic for selecting an multi-plexing the on a limited number of I/O pins. The chip has beenimplemented with standard CMOS technology (with the addi-tion of gold deposition for capacitor electrodes). Capacitancemeasurements are performed by means of the charge-based ca-pacitance measurement (CBCM) technique [23], which is suit-able for low-complexity, small-area on-chip implementation.

The chip includes addressing circuitry, while the read-outelectronics has been realized externally by means of standardcomponents in order to exploit the advantages of large exper-imental flexibility (although it could eventually be easily inte-grated with the rest of the system).

III. CAPACITANCE-BASED DNA DETECTION PRINCIPLE

Under suitable electrochemical conditions, bio-modifiedmetal interfaces in saline solution exhibit a capacitive behavior.This is the case if gold electrodes are modified with short DNAstrands immobilized with alkanethiol linkers on the surface[18], [20]. In this case, capacitance values of electrode/solu-tion interface have been estimated between 1 and 20 F cm(although this value strongly depends on electrode surfacetreatment). A first-approximation model of the electrode/solu-tion interface is the equivalent circuit depicted in Fig. 1 (left)

Fig. 1. Left: schematic representation of the lumped-element electrical modelof the metal–solution interface of the capacitors used in this work. Right:schematic illustration of the DNA hybridization process and the induced(further) displacement of counterions within the liquid solution.

Fig. 2. Schematic representation of the system and signal flow. A and Dindicate analog and digital signals, respectively.

[24], [25], where depends on the interface and solutioncharacteristics; is related to the insulating properties ofthe interface (for well-formed layers it can be consideredinfinite); is mainly affected by the physical and chemicalcharacteristics of the insulating biolayer immobilized on thesurface. In this model, the geometrical capacitance formed bythe electrodes would be in parallel to but, since it is severalorders of magnitude smaller than the interface-capacitance, itscontribution is negligible: thus, it is not considered here.

This simple model provides an intuitive explanation of thebasic sensing principle exploited in our work. Considering a ca-pacitor formed by two neighboring electrodes of our chip, whena complementary DNA strand binds with the surface probes andthe DNA duplex is formed, the solution counterions present inthe liquid solution are pushed further away from the polarizedmetal surface: this increases the distance between the charge in-side the electrodes and the ions in the electrolyte: thus, de-creases [19]–[21] (Fig. 1, right).

When measuring a capacitance with CBCM technique, aninput square wave is applied to one electrode, while the currentrequired to completely charge (or discharge) the capacitance ismeasured at the second electrode. The measured capacitancevalue equals the product of the average value of the output cur-rent, the applied voltage step, and the frequency of applied clocksignal. With this technique, the measurement of the capacitivecomponent of an impedance is independent from the resistive

STAGNI et al.: FULLY ELECTRONIC LABEL-FREE DNA SENSOR CHIP 579

Fig. 3. Schematic of the circuit associated to each sensing element.

components provided that 1) the charging and discharging cur-rent transients have the time to settle during the period of thesquare wave (i.e., all the charge stored in the reactive compo-nent is transferred) and 2) the static current flowing through the

and resistive path is negligible compared with the av-erage output current (i.e., most of the charge transfer is due tothe reactive component). Previous work on CBCM has demon-strated that both these requirements can be met in a practicalexperimental setting [21].

It is important to stress, however, that the model of Fig. 1 isa highly idealized view of more complex physical phenomena,namely charge transfer at the electrode–solution interface inpresence of biological molecules immobilized on the electrodesurface. Thus, we cannot assume, for instance, that capacitancesand resistances in Fig. 1 are constant over a wide frequencyrange. Sensing of impedance variations requires, therefore,careful selection of the frequencies used for CBCM excitationwaveforms. These issues are addressed in Section VI.

IV. CHIP

This section describes the architecture and the operation ofthe test chip designed, fabricated, and characterized in this work.The chip contains a number of sensing sites and the addressingcircuitry, while read-out is realized by means of by externalcomponents in order to systematically explore different mea-surement configurations and parameters. The chip is packagedon a PCB to form a system suitable for test measurements andcharacterization.

A. Chip Architecture

Essentials of chip architecture and signal flow are shown inFig. 2. Both analog and digital input signals are required. Theformer ones determine the voltage range used for the measure-ments, while the latter ones include a seven bit address to selectthe sensing sites. An external clock is used to provide the mea-surement frequency, which is the same as the external signal,as well as to synchronize the different blocks. Each sensing sitefeatures a block that, controlled by external clock, generates afour square-wave signal with high timing precision.

Fig. 4. Schematic representation of the signals flow used in the experiments.

The on-chip address signals are generated by means of twodecoders. If one sensing site is selected, the input signals comingfrom the pulse generator are enabled and measurement is per-formed. Only the output signal of the selected site is connectedto the shared output pad (Fig. 2).

At the output pad, the transient current is converted by meansof an external circuit into a voltage signal, that is sampled andprocessed by a PC to calculate the capacitance value. With re-spect to an on-chip implementation of the – conversion, thisexternal solution allows larger flexibility in experimenting dif-ferent measurement parameters.

B. Sensing Site Circuitry

The circuit implemented for each sensing sites, illustrated inthe dotted box in Fig. 3, includes a pulse generator and fourswitches which selectively connect the electrodes of the sensingcapacitances to one of three different reference voltages: ,

, and . Between the pulse generator and the switchesa multiplexer is inserted which allows the clock pulses to drivethe electrodes of the sensing element only if it is selected. Theswitches are implemented with both n- and p-channel transistorsin order to exploit full voltage range.

Exploiting its proximity with the switches, the (local) pulsegenerators provide very precise signals controlled by the ex-ternal clock. A critical timing issue has to be considered at this


Fig. 5. Schematic plot of the block used to generate not overlapping clock signals.

TABLE IDESCRIPTION OF THE CHIP PINOUT

regard in that , , and , must not close theswitches at the same time in order to avoid shorting and

. For this reason, a suitable delay is inserted and nonover-lapping clock signals are used (as shown in the inset of Fig. 4)in order to avoid that two transistors at the same time see a gatevoltage larger than their threshold. The delay used for this pur-pose is 1 ns and is obtained by means of the pulse generatorblock reported in Fig. 5, also showing the output buffer fol-lowing the multiplexer.

The switches allow to connect the electrodes of the sensingcapacitors either to and to , or to and the voltageabout at the negative input of the opamp shown in Fig. 5(see also Table I). When switching between these two bias con-ditions, a voltage step is applied at the sensing electrodes and atransient current is generated. With the circuits shown in Fig. 3,the charging and discharging currents are driven to two differentoutput pads, OUT1 and OUT2, thus allowing measurement ofthe average value using only one of them.

C. Physical Layout

The chip is fabricated in 6-in n-well 0.5- m CMOS processwith two metal layers. The gold electrodes are deposited after

Fig. 6. Chip photograph.

standard CMOS processing on silicon nitride (Si N ) as de-scribed in [4]. Annealing steps are introduced to guarantee goodCMOS performance after gold deposition [4]. The area of thethree sensing capacitors used in this work is 1 mm 1 mm andthe distance between them is 500 m, while that of total chip,featuring 44 electrode couples of different size, is 6.5 mm 4.6mm (Fig. 6). The parasitic capacitance between gold electrodesand substrate has been estimated to be 6 aF m , therefore, itis several order of magnitude smaller than electrode/solution in-terface capacitance (about 35 nF as shown in Section VI).

The pin-out includes analog as well as digital signals, powersupply and ground. Table I provides all I/O signals, with indi-cations of the label and the value of the voltage applied duringthe measurements.

The main analog signals are , , determining thevoltage step ) applied to the electrodes.Digital inputs include the seven bit address signals and oneexternal clock to set the measurement frequency. The powersupply voltage is the same for the whole chip.

In order to reduce interferences and noise, analog ground is. Our choice of and allows to gen-

erate only positive signals (e.g., , ), while the digitalground can be made negative respect to . In fact, the signalvoltage levels , and

are generated by external voltage sources relatedto digital ground.


Fig. 7. Photo of the PCB used for measurements. The chip is highlighted in thebox on the right.

Fig. 8. Schematic representation of the measurement setup.

V. MEASUREMENT SETUP

The chip described in the previous section is glued to a PCBfeaturing large contacts on one side to be connected to standardinstrument and small ones on the other side connected to thechip by means of gold bonding wires (Fig. 7). A passivationprocess protects both wires and chip from the saline solutionsused in the experiments. To this purpose, epoxy glue is placedon the critical electrical parts. Since to bind the DNA probesto the gold electrodes, these must be accurately cleaned withplasma oxygen and epoxy glue is damaged by this process afurther SU-8 layer is added to protect the glue from treatment.

In our electrical measurements, the transient discharging cur-rent is driven to an external – converter (Fig. 8) implementedoff-chip by means of the operational amplifier OPA103AM[26] and precision resistors. The voltage signal at the op-ampoutput is sampled by means of a National Instruments DAQboard (PCI-6032E [26]). Finally, the average value of theoutput voltage ( ) is calculated by means of a trapezoidalintegration and the average discharge current ( ) is thenobtained as

(1)

can be expressed as

(2)

Here, , , and are controlled parameters,while is the capacitance to be extracted from experimentaldata.

For higher precision, the value of is measured at ninedifferent frequencies (from 60 to 100 Hz and step of 5 Hz foreach measurement) as the average of 10 charging transient.

Finally, is computed by least squares fitting on (2) pro-viding the slope

(3)

VI. EXPERIMENTAL RESULTS

A. Electrical Characterization of Gold Electrodes

Our chips have been tested in wet conditions at room temper-ature. After cleaning, a saline solution, TE (Tris 10 mM, EDTA1 mM) 0,3 M NaCl pH 7, is spread on the electrode surfaceand CBCM measurements are performed to characterize differ-ences between electrode couples. In a first phase, test measure-ments are performed to set two important parameters, namely

, which determines the – conversion factor, and the fre-quency range to be used in the measurements.

From Section III, we recall that CBCM frequency should below enough to reach steady state in cell current, and that the dccurrent flowing through the and resistive path should benegligible compared with the average output current. By ana-lyzing current transients in response to voltage steps, we deter-mined that steady state was reached in a few tenths of microsec-onds. We, therefore, set the period of the excitation waveformone order of magnitude larger than transient time. Moreover,and have been estimated with a standard impedance spec-troscopy measurement and found to be about 10 M and 10 ,respectively. is large enough to ensure that the dc currentflowing between the electrodes falls below the noise level of theexperimental setup. Thus, both requirements for applicability ofCBCM are met.1

However, the experimental measurements of , in contrastwith the first-approximation model of Fig. 1, (left), have showna large frequency dependence (Fig. 9). The behavior can beattributed to electro-chemical effects taking place at the elec-trode-solution interface of the sensing device. These effects can,in fact, be accounted by modeling with a constant phase el-ement [27]. Indeed, Fig. 9 reports a good fit of our experimentaldata obtained with the described model.

It is worth pointing out that detection does not require deter-mining a single value of the sensing capacitance, but to distin-guish between the device behavior in the presence rather thanin the absence of DNA hybridization and, as shown later, this isindeed the case.

The – converter is implemented with the operational am-plifier OPA103AM [26] featuring a gain-bandwidth product of10 , a value of interest in the perspective of fully integratedelectronics. This characteristic plays a crucial role in the choiceof the frequency range. The optimal value of for the CBCMvoltage excitation frequency to be used in the measurements is

1To further strengthen this point, PSPICE simulations have been done con-sidering the circuit represented in Fig. 1, (left), with a constant value of C .With this model, the capacitance is then extracted from the (average) transientdischarge current as done in the experiments (described in Section V) and theresults indicate a dependence on f smaller than 7% in the frequency range from10 Hz to 1 kHz).


Fig. 9. Measured capacitance versus charge/discharge frequency on cleangold electrodes. The continuous line shows the fitting.

chosen taking into account the frequency dependence shown inFig. 9 and also the characteristics of the – converter. Lowerfrequencies (hence, longer periods) lead to long measurements,with lots of noise. On the other hand, when frequency is toohigh, small gains of the – converter lead to small signals,again critically affected by noise.

In practice, if is 1 kHz, the gain-bandwidth product of ouramplifier would require k . Unfortunately, in this case,gain would be rather low, and the signal-to-noise ratio (S/N)would be too low (about 5 dB). On the other hand, ifHz, then k and gain would be high enough, but themeasurements would be too noisy, as shown in by the error barsof Fig. 9.

Therefore, the optimal condition of the – converter circuitis k (leading to dB and total measure-ment time around 20 s for each site), and in the range from 10to 100 Hz. The frequency dependence of suggests to workat a measurement frequency as high as possible. Therefore, wehave used frequencies in the range from 60 to 100 Hz.

Typical experimental results, obtained with (3) are shown inFig. 10 and Table II. “Couple 1” refers to the electrode plateson the left of the chip while “Couple 2” refers to those on theright. The standard deviation reported in Fig. 10 come from sev-eral measurements for each sensing site. Relative differences be-tween capacitance measurements for two different couples onthe same chip are evaluated to be less than 4.2%. Moreover, sta-tistics on data shows that the capacitance values measured onthe same chip but on two different couples are almost the same.Differences among several chip are due mainly to differences inthe level of cleaness of the electrodes after the cleaning step.

B. Electrode Biomodification

First, the gold electrodes have been cleaned by exposurefor 20 min to plasma oxygen at 200 W. After this step, singlestranded DNA molecules modified with alkanethiol groups areimmobilized on the gold electrodes by covalent S–Au bonds (a3- M DNA 1 M Na HPO solution is spread on the electrodesurface for 18 h). Two different probe molecules of the samelength (25-mer) and thiol modified with a chain of six carbonatoms as a spacer are bound to different electrodes on the same

Fig. 10. Capacitance measurements of electrode couples on different chips.

TABLE IICAPACITANCE MEASUREMENT ON SIX DIFFERENT CHIPS

chip (electrode couple on the left and on the right). To this end,two separate droplets are placed by mean of microliter pipettesto obtain sites which will and will not experience hybridizationreaction. Before measurements, the gold surfaces are exten-sively rinsed with ultra pure water to remove molecules that arenot covalently bound to the gold electrodes or to the passivationlayer of the chip.

Target DNA solution (3- M DNA 30-mer and TE 0,3 MNaCl pH 7) is heated up to 80 C, spread on the electrodes andcooled down to room temperature (for about 30 min). Finallythe sample is rinsed in the same saline solution (TE 0,3 M NaClpH 7) in order to remove the unbound DNA target (Fig. 11).

In order to verify the biological steps previously described,we have performed an independent standard optical detec-tion test based on fluorescence molecules bound to DNAmolecules. Probe molecules are labeled with fluorescein andtarget molecules with tetrametil-rhodamine. In order to performquantitative analysis after electrical measurements the DNAmolecules, both probes and targets have been removed fromthe surface by mean of -mercaptoethanol 12 mM solutionfor 10 hours. Then, to increase the optical signal, the pH ofthe sample solution has been brought to 10 adding a solutionNaOH 10 M (the pH has been checked with the HI-9023pH meter). Before DNA experiment a calibration curve hasbeen done with the LSB-50 fluorescence spectrometer (PerkinElmer, Boston, USA) using DNA sample solutions, (12 mM

-mercaptoethanol and pH 10 adding NaOH 10 M solutionas previously described) of different known concentrations.


Fig. 11. Schematic representation of the functionalization and hybridizationsteps.

Results of optical detection of our unknown samples com-pared with the calibration curve show that the density of ourprobes layer is about 10 molecules cm , a value close to thatreported in the literature [28]. Moreover, we have tested theefficiency of the hybridization reaction in the case of comple-mentary and a-specific target molecules: the first one indicatesthat the 80% of the probes react with target forming the doublehelix while the second one indicates that less than 10% of probereact with target, demonstrating the good quality of our process.

C. DNA Detection

DNA detection is demonstrated comparing measurements onelectrode couples subjected to the same reaction but with dif-ferent DNA strands bound on the surface, complementary andnon complementary to target molecules, respectively (the latterfor negative control). All measurements are performed in thesame saline solution of the hybridization step (TE 1X 0,3 MNaCl pH 7). Since capacitances exhibit significant mismatches,a measurement after functionalization is performed and thesevalues are used as a reference to be compared with the resultsobtained after (tentative) hybridization. The capacitance abso-lute values after functionalization show a decrease around 25%respect to clean gold for each couple. Results of three experi-ments over all the chip previously measured after cleaning aresummarized in Fig. 12, where positive variations correspond tocapacitance decreases.

Fig. 12 clearly indicates that significant capacitance decreasesare found as a result of hybridization, while smaller or even op-posite variations occur in the other case, thus demonstrating thecapability of our approach to distinguish between hybridizationand nonhybridization events.

Fig. 12. Capacitance variations due to specific and a-specific bindings (upperand lower bands of measured capacitances, respectively). Positive valuesindicate capacitance decrease.

Fig. 13. Capacitance variations due to dehybridization process in the case ofcomplementary (lower) and noncomplementary (higher) probes previouslysubjected to hybridization reaction. Negative values indicate capacitanceincrease.

After measurements, the sensing layers have been regener-ated removing all target molecules. For this purpose, a rinsingstep with hot ultrapure water is performed to break hydrogenbonds between the two DNA strands, leaving only the probeoligonucleotide chains covalently bound on the surface. Afterthis step, new capacitive measurements are performed andFig. 13 illustrates the measured capacitance variations.

As expected, an increase of interface capacitance is observedwhere hybridization had previously occurred, while electrodeswith noncomplementary probes show only negligible variations.Moreover, it is worthwhile to note that the absolute variationsof capacitance in dehybridized chips are of the same amountoccurring in hybridized ones.

VII. CONCLUSION

In view of a simple and effective DNA analysis, a significanteffort is currently dedicated to the development of single-chip,fully electronic devices and this paper illustrates a step forwardin this direction since it demonstrates the viability of a simpleapproach suitable for stand alone and portable diagnosis equip-ments based on two-electrode capacitive measurements imple-mented with microfabricated capacitors.

In particular, this work presents a test chip, fabricated withstandard CMOS technology with the addition of gold depositionfor the sensing capacitors. The chip includes the addressing cir-cuitry, while standard external components are used to realize


the read-out electronics, in order to improve measurement flex-ibility and help the search for optimum conditions.

The chip has been fully characterized and measurements havebeen performed exposing the device to target DNA solutions.The results clearly indicate that DNA hybridization produces asignificant decrease in the capacitance of the sensing capacitors,that can be safely recognized by the on-chip read-out circuitry.

The paper then suggests that a chip, potentially containingthousands of sensing sites, with fully integrated circuitry forDNA recognition can be fabricated exploiting the approach de-scribed in this work.

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Claudio Stagni was born in Pavullo, Italy, on De-cember 17, 1977. He received the B.S. degree in elec-trical engineering from the University of Bologna,Bologna, Italy, with a thesis ”Capacitance measure-ments to detect DNA hybridization, ” in 2002, wherehe is currently working toward the Ph.D. degree indesign and test of low cost biosensors.

In 2004, he was with Infineon Technologies,Munich, Germany, where he worked in the field ofCMOS DNA sensors.

Carlotta Guiducci was born in Bologna, Italy, in1977. She received the B.S. (summa cum laude)and Ph.D. degrees in electrical engineering from theUniversity of Bologna, Bologna, Italy, in 2001 and2005, respectively.

Since 2002, she has been with the MicroelectronicGroup and with the Nanobioscience Group, Univer-sity of Bologna, working on the development of inno-vative DNA and protein sensors for point-of-care ap-plications. In 2003, she was a Visiting Researcher atthe LMGP-INPGrenoble where she investigated the

employment of conductive transparent oxides for electrical DNA detection. Sheis now a Visiting Professor at the ESPCI-Paris, France, where she works onfield-effect devices for molecular analysis based on charge characterization.


Luca Benini (S’94–M’97–SM’04–F’07) receivedthe Ph.D. degree in electrical engineering fromStanford University, Stanford, CA, in 1997.

He is a Full Professor at the Department of Elec-trical Engineering and Computer Science (DEIS),University of Bologna, Bologna, Italy. His researchinterests are in the design of system-on-chip plat-forms for embedded applications. He is also activein the area of energy-efficient smart sensors andsensor networks. He has published more than 250papers in peer-reviewed international journals and

conferences, four books, and several book chapters. He has been ProgramChair and Vice-Chair of Design Automation and Test in Europe Conference.He has been a member of the Technical Program Committee and OrganizingCommittee of several technical conferences, including the Design AutomationConference, International Symposium on Low Power Design, the Symposiumon Hardware-Software Codesign.

Dr. Benini is an Associate Editor of the IEEE TRANSACTIONS ON COMPUTER-AIDED DESIGN and the ACM Journal on Emerging Technologies in ComputingSystems.

Bruno Riccò (F’03) was born in Parma, Italy, on Feb-ruary 8, 1947. In 1971, he graduated in electrical en-gineering from the University of Bologna, Bologna,Italy, and in 1975 received the Ph.D. degree from theUniversity of Cambridge, Cambridge, U.K., where heworked at the Cavendish Laboratory.

In 1980, he became Full Professor of AppliedElectronics, University of Padua, Padua, Italy, andin 1983, he joined the University of Bologna. Hehas been Visiting Professor at the University ofStanford, at the IBM Thomas J. Watson Research

Center, Yorktown Heights, NY, and at the University of Washington. In 1999,he was appointed European representative for the International ElectronDevice Meeting (IEDM). He has been consulting for major semiconductorcompanies and for the Commission of the European Union. From a scientificpoint of view, he has worked in the field of solid-state devices and ICs. He is(co-)author of over 300 publications, more than half of which have publishedon major international journals, of three books, and of six patents in the fieldof nonvolatile memories.

Dr. Riccó received the G. Marconi Award by the Italian Association of Elec-trical and Electronics Engineers (AEI), for his research in electronics in 1995.

Sandro Carrara graduated in physics from the Uni-versity of Genoa, Genoa, Italy, in 1990 and receivedthe Ph.D. degree in biochemistry and biophysics atthe University of Padoa, Padoa, Italy, in 1998.

He began his research activity in the field of theSTM microscopy. During the year he spent in thePh.D. program, his interest was on the electronconductivity phenomena on thin films and he wona prize in a NATO Advanced Research Workshopfor his work on the role of nanoparticles size inmono-electron conductivity, in 1996. He was Assis-

tant Professor in Biophysics. He is currently a Research Associate with theDepartment of Biochemistry, University of Bologna, Bologna, Italy.

Christian Paulus received the diploma degree inphysics from the University of Bayreuth, Bayreuth,Germany, in 1997.

From 1997 to 2005, he has been with and InfineonTechnologies AG, Corporate Research, Munich,Germany, and Siemens AG, Corporate Technology,Munich, where is he currently working. There heworked on the development of cell based and molec-ular biosensors as well as on high-speed analog todigital converters. He holds more than 50 patentsand is author and coauthor of more than 30 national

and international publications.

Meinrad Schienle received the Dipl.Phys. and Ph.D.degrees from the University of Karlsruhe, Karlsruhe,Germany, in 1979 and 1983, respectively.

In 1984, he joined the Research Laboratoriesof Siemens AG, Munich, Germany, working onintegrated optical components based on III–V semi-conductors. Since 1999, he has been with CorporateResearch, Infineon Technologies AG, working ondesign-based yield evaluation and circuit design forflash memories and the development of fully elec-tronic DNA sensors. Since 2005, he has been with

Siemens AG, Corporate Technology, working in integrated CMOS sensors.

Roland Thewes was born in Marl, Germany, in 1962.He received the Dipl.-Ing. degree and the Dr.-Ing. de-gree in electrical engineering from the University ofDortmund, Dortmund, Germany, in 1990 and 1995,respectively.

From 1990 to 1995, he worked in a cooperativeprogram between the Siemens Research Labora-tories, Munich, Germany, and the University ofDortmund in the field of hot-carrier degradationin analog CMOS circuits. In 1994 , he joined theResearch Laboratories of Siemens AG and Infineon

Technologies. From 2000 to 2005, he was responsible for the laboratory onMixed-Signal Circuits of Corporate Research of Infineon Technologies fo-cusing on CMOS-based bio-sensors, device physics-related circuit design, andadvanced analog CMOS circuit design. Since 2006, he is heading a departmentdeveloping DRAM core circuitry at Qimonda AG, Neubiberg, Germany. Hehas authored or coauthored approximately 120 publications including bookchapters, tutorials, invited papers, etc. He served as a member of the TechnicalProgram Committees of the International Reliability Physics Symposium(IRPS), and of the European Symposium on Reliability of Electron Devices,Failure Physics and Analysis (ESREF). He is a member of the TechnicaLProgram Committees of the International Solid-State Circuits Conference(ISSCC), of the International Electron Device Meeting (IEDM), and of theEuropean Solid State Device Research Conference (ESSDERC).

A Fully Electronic Label-Free DNA Sensor Chip

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