Sensing Cellular Adhesion with a CMOS IntegratedImpedance-to-Frequency Converter

Andreas Mucha, Meinrad SchienleCorporate Research and Technologies

Siemens AGMunich, Germany

Email: andreas.mucha@mytum.de

Doris Schmitt-LandsiedelLehrstuhl fur Technische Elektronik

Technische Universitat MunchenMunich, GermanyEmail: dsl@tum.de

Abstract—Sensing cellular adhesion via impedance measure-ments provides a versatile and easily accessible means formonitoring in-vitro cell cultures. In this work we present a CMOSimpedance-to-frequency converter integrated with biocompatibleplanar surface electrodes to make a compact and robust sensorchip for in-vitro cell monitoring. The system features an 8×8array of individually addressable electrodes connected to fourimpedance-to-frequency converter circuits with square waveoutput. The impedance-to-frequency converter circuits consistof a switchable constant current source, two comparators, andassociated logic for addressing the pixel array and generatingcontrol signals. In designing the electrodes and circuit compo-nents, special attention had to be paid to keeping the currentdensity and polarization voltage at the electrode low enough toavoid damaging the cells or triggering unwanted electrochemicalreactions. To this end, FEM simulation was used to optimizeelectrode layout and electrical circuit parameters, keeping theelectrode current below 100 nA. We present measurement resultswith cells that demonstrate the successful operation of the systemand show good agreement with models of the electrode and cellimpedances.

I. INTRODUCTION

Real time monitoring of live cells in lab-on-chip systems hasproven useful in various applications [1]: In pharmaceuticalresearch, it allows e.g. an effective screening of new drugcandidates, thus reducing costs and the number of necessaryanimal experiments. Another promising application is thedetection of environmental toxins, where mammal or evenhuman-derived cell lines make highly sensitive yet broad-band indicators for substances that could harm man or animal.Signals from such a cell based sensor have a high intrinsicvalidity, because anything that affects cultured cells also has ahigh probability of affecting the whole organism. Their non-selectivity sets them apart from classical chemical analysismethods, which can achieve high sensitivity, but are alsohighly specific, which means that only substances that aretested for can be found. Living cells on the other hand reactand adapt to a multitude of substances, e.g. by changing theirrate of proliferation, morphology, and adhesion to a substrate.These reactions are detectable by measuring impedance, aswill be explored in more detail in section II-B.

Advances in micro-fabrication, mainly driven by integratedcircuit technology, today allow the integration of various typesof sensors together with signal processing and interface elec-

tronics on the same die. The advantages of such an integrated”smart” sensor are numerous: Firstly, due to various factors,the signals generated by the sensors are generally quite weak,i.e. in the µV or nA range. The reduced physical distancebetween the transducer element and the associated front-endelectronics translates into less susceptibility to interference anda higher signal to noise ratio. Secondly, with the integrationof signal processing, multiplexing, and analog-to-digital con-verter circuits on the sensor chip, the number of requiredconnections is reduced accordingly. This also improves therobustness of the measurement setup. Thirdly, CMOS tech-nology as the de facto standard for integrated circuits offersthe opportunity to produce an inexpensive integrated cell basedsensor chip, thus for the first time making such a biosensorsuitable for widespread application.

Previously published cell impedance sensors [2–4] used ex-ternal electronics wired to microelectrodes, and thus incurredparasitic impedance, electromagnetic interference, and bulkymeasurement setups, negating key advantages of lab-on-chipsystems. In this paper, we present results of the developmentof an integrated CMOS impedance-to-frequency converter forsensing cellular adhesion, intended for application as part ofa multi-parameter cell sensor for water quality monitoring.

II. THE IMPEDANCE OF CELLS ON PLANAR ELECTRODES

The first step towards designing an impedance sensor forcell adhesion is the development of a suitable model of theimpedance to be measured. This model can be broken downinto three parts: The first part describes the impedances of theelectrode and the electrolyte. These can be suitably describedby their respective resistances, which are small in relation tothe other impedances in the system at the relevant frequencies.

The electrode-electrolyte interface however requires a moredetailed analysis, described in the following section. The thirdpart of the model describes the impedance arising from thecells on the sensor and is also described in its own section.

A. Impedance of the electrode-electrolyte interface

The impedance of the electrode-electrolyte interface hasbeen a subject of study in electrochemistry for a long time,and consequently there is no lack of literature concerningthis topic. Fig. 1 shows one of the equivalent circuit models

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Fig. 1. Assumed equivalent circuit diagram of a metal electrode in contactwith an electrolyte. The polarization impedance Zp is infinite at 0Hz, howeverthe charge transfer resistance Rct allows direct current across the electrode-electrolyte interface if a high enough voltage is applied. The spreadingresistance through the electrolyte is represented by Rel

that are commonly used to describe its behavior. An impor-tant distinction to make here is whether an actual transferof charge carriers across the electrode-electrolyte interfaceoccurs or not [5, 6]. If a voltage is applied to a pair ofelectrodes connected by an electrolyte, oxidation reactionsat the anode (+) and reduction reactions at the cathode (-)occur. The magnitude of the applied voltage together withthe activities of the involved ions determine the faradaic orcharge transfer resistance Rct, which arises from the passageof charge carriers through the electrode-electrolyte interface.This is accompanied by changes in the electrode surface, asreaction products adsorb to the electrode or metal ions gointo solution. As this would preclude measurements of anysignificant duration, it is important for the operation of theimpedance sensor that the applied excitation is small (voltagesof tens of millivolts and current densities of few milliamperesper square centimeter), so that Rct can be considered infinite.

In this case, only the polarization impedance of theelectrode-electrolyte interface needs to be considered. How-ever, most published models of the polarization impedance of(solid) metal electrodes are phenomenological in nature, asthere is still no consensus on its origin and the mechanismsinvolved [5, 7, 8]. In our experiments, we found that the modelpublished by Ragheb and Geddes [9] resulted in the best fitto the measured data. It takes the form of

Zp = Af−α − jBf−β (1)

with the complex impedance Zp in ohms and A, B, α and βpositive constants.

Preliminary impedance spectroscopy measurements per-formed on gold electrodes suspended in cell culture mediumyielded average values A = 1.09 × 103, B = 9.54 × 103,α = 0.99, and β = 0.91 (normalized to 1 cm2). Since B > Aand more importantly α > β, the real part of Zp becomesnegligible at higher frequencies and the electrode polarizationimpedance can be approximated as a frequency-dependentcapacitance, with a value around 10 µF cm−2 below 100 kHz.

B. Modeling the impedance of a cell monolayer

The measured impedance of adherent cells growing onplanar electrodes correlates with the electrode area coveredby the cells, i.e. with the cell number [3], and in the caseof monolayer-forming cells also with the tightness of thejunctions between adjacent cells [10].

Fig. 2. Cross section of the geometry assumed in the model [12] and our FEManalyses: A current I0 is injected into a planar electrode with polarizationimpedance Zp on the left and spreads in the space of height h under a cellmonolayer, represented by a lumped impedance Zcells, and between the cellsinto the bulk medium, which is assumed to be at a fixed potential Vm. Thedashed line at the left indicates the axis of symmetry.

However, attempts to capture the behavior of the cell layerimpedance in a comprehensive analytical model have not beencompletely satisfying so far, often relying on non-obviousassumptions [10–12]. The model developed by Urdapilletaet al. [12] seemed most promising since it does not assumea particular shape of the cells, accounts for a finite electrodesize and seems to fit experimental data best. It is howeverlimited to radially symmetric electrode geometries since noanalytical solution for the involved differential equations existsin the two-dimensional case. We therefore combined theirmodel with numerical analyses using ANSYS finite elementssoftware to calculate the expected magnitude of the cell layerimpedance in the case of other electrode geometries.

Fig. 2 shows a cross-section of the geometry which the an-alytical model by Urdapilleta et al. [12] and our own finite el-ements analyses were based on. The underlying assumption isthat the current injected through the electrode is forced throughthe gaps between or under the cells, creating a significant con-striction resistance, because the cell membrane has a very highimpedance compared to the resistivity of physiological media.The impedance of the cell monolayer is modeled as a parallelRC circuit, using published values of 0.3 kΩ cm2 to 2 kΩ cm2

and 1 µF cm−2 to 4 µF cm−2 [2, 10, 13]. The resistivity of thecell culture and assay medium (Dulbecco’s Modified EagleMedium) as measured by ourselves is 65 Ω cm.

A critical parameter in determining the total impedanceis the height h of the cell-substrate gap, since it directlyinfluences the constriction resistance of the cell layer. A highervalue of h means less current constriction and thus lowersignals. In accordance with published values [12], we chose aworst-case value of 2 µm for our FEM simulation. To ensurethat the cells are not affected by the measurement, whichwould falsify the results, a current density of about 2 mA cm−2

at the electrode should not be exceeded. As a result of theseconsiderations, we chose a layout of square electrodes of55 µm edge length. According to our calculations, this shouldgenerate a resistive cell layer impedance of around 130 kΩwhile keeping the allowed electrode current at a manageablelevel.

III. IMPEDANCE-TO-FREQUENCY CONVERTER DESIGN

The design of our impedance-to-frequency converter isbased on the results of the modeling described in the pre-

2I0

I0Rel

Cel

Φ1

Φ2

VrefL

VrefHVel Vout

Fig. 3. Simplified block diagram demonstrating the principle of operationof the impedance-to-frequency converter

vious section: Its measurement principle exploits the quasi-capacitive behaviour of the electrode-electrolyte interface atlow polarization voltages in conjunction with the insulatingproperties of biological cell membranes by injecting a constantcurrent in the nA range into the electrode while keeping thebulk electrolyte at a fixed potential. A simplified block diagramof the circuit is depicted in Fig. 3. While the output of thecomparator Vout is high, Φ1 is active, Φ2 is low and thearrangement of the two current sources results in a currentI0 injected into the electrode. The comparator continuouslymeasures the electrode potential Vel and switches its outputVout to low as soon as the upper threshold VrefH is reached.A non-overlapping pulse generator ensures that Φ1 goes lowbefore Φ2 becomes active, as a short-circuit to ground wouldexpose the cells to an unwanted high current. In the resultingswitch configuration, the electrode is discharged by −I0. Thepurpose of the switch controlled by Φ2 is to keep the channelof the MOSFET in the upper current source active, so thatit can instantly resume operation as soon as Vel reaches thelower threshold VrefL and Vout switches again, completingthe cycle.

The time needed to charge and subsequently discharge theelectrode potential between the two reference voltages asillustrated in Fig. 4 determines the frequency of the outputsignal. As the current through the resistance Rel changesfrom +I0 to −I0 (or vice versa), an ohmic voltage drop ofmagnitude 2RelI0 occurs which effectively reduces the voltagewindow of ∆V = VrefH−VrefL. Neglecting parasitic effects,the period of one charge-discharge cycle is thus given by

T = 2CelI0

(∆V − 2RelI0) = 2Cel

(∆V

I0− 2Rel

)(2)

A higher resistance Rel caused by cell attachment thus trans-lates into a lower value of T and a higher frequency of Vout.

The values for ∆V and I0 have to be chosen as a com-promise between the constraints imposed by the cells, which

ΔV

2RelI0

V

VrefL2RelI0

t

Tel

VrefH

V

t

out

Fig. 4. Idealized plot of the electrode voltage and digital output over time

demand low values for both, and the necessity to separatethe cell layer impedance from the background signal of theelectrode-electrolyte interface impedance, i.e. to obtain a goodSNR. The latter can be achieved by using sufficiently highexcitation frequencies where the polarization impedance issmall, which means higher values of I0. The limitations ofthe circuit implementation by noise and leakage currents alsohave to be considered. With these factors in mind, we designedour sensor for a value of ∆V = 20 mV and the current I0adjustable in a range of 10 nA to 100 nA. This was expectedto result in frequencies around 10 kHz.

The output of the comparator provides a digital frequency-modulated output signal which is easily processable. Celladhesion is detected as a rise of the output frequency, whicharises when the high spreading resistance of a cell-coveredelectrode causes an ohmic voltage drop. Thus, while the outputof our impedance-to-frequency converter does not exactlycorrespond to an impedance magnitude, the output frequencycan be interpreted as a cell adhesion index that contains allthe useful information for the user.

IV. MATERIALS AND METHODS

A test chip for the impedance-to-frequency converter circuitwas fabricated in a standard 0.35 µm 3.3 V single-well CMOStechnology supplemented with a backend process for planargold electrodes similar to the one described in [14]. The testchip features 64 sensing sites for cellular adhesion, arranged inarrays of four by four. Each of the 64 electrodes is individuallyaddressable by a digital interface, with the four electrodesin the center also accessible off-chip via a pad for testingand control measurements. Two counter electrodes with anarea much larger than the sensing electrodes maintain theelectrolyte at an externally-supplied fixed potential.

The chips were housed in an open CPGA cavity pack-age (Spectrum Semiconductor Materials) and the bond wiressealed with Vitralit 1671 (Panacol). A polystyrene dish 35 mmin diameter with a hole corresponding to the package cavitycut out in the center was then glued tightly on top, also withVitralit 1671, to form a liquid reservoir.

For the experiments with cells, V-79 cells were cultured inan atmosphere of 5 % CO2 at 37 C in Dulbecco’s Modified

Fig. 5. Photograph of a packaged chip mounted on the support PCB

Eagle Medium supplemented with 10 % FCS. During themeasurements, the chip and support PCB were placed in theincubator as well to maintain these conditions. The liquidreservoir was filled to contain 2 ml of medium. The biasing ofthe chip was adjusted for an electrode current of 70 nA. Forcomparison of the output of our impedance-to-frequency con-verter to actual impedance, we connected a Hewlett PackardHP4284A LCR meter to the chip. This allowed us to performimpedance spectroscopy on the four center electrodes of thechip and at the same time measure the remaining 60 electrodeswith the on-chip circuitry. The excitation voltage of the LCRmeter was set to 20 mVrms with no DC bias.

A multifunction I/O box attached to a PC was used to outputthe digital address bits and to supply power to the chip. Atimer in the same I/O box was used to evaluate the frequency-modulated digital output of the chip.

V. RESULTS AND DISCUSSION

A. Detection of cell reactions

We seeded cells on our chip and let them grow for about22 hours until a confluent monolayer had formed and theimpedance signals had stabilized (data not shown). Then, weexchanged 1 ml of the used medium with fresh medium. Inthe graph shown in Fig. 6, a short dip of the chip outputfrequencies as well as the impedance measured by the LCRmeter can be seen immediately after the exchange, followed bya rise to a higher level than before. Since the fresh medium wasat room temperature and the experimental environment waskept at 37 C, the dip could be due to the temperature shock,while the new nutrients and higher oxygen content explain therise of the impedance afterwards. It is important to note that inthe data from the control experiment performed without cells(dashed lines), the exactly equal treatment results in almost noeffect.

After the signal had stabilized again, we exchanged 1 ml ofthe medium on the chip with 1 ml of medium containing 0.2 %of Triton-X 100, a detergent that kills cells by destroying theirmembrane. An immediate drop of the impedance down to the

1 ml of freshmedium added

1 ml of0,2% Triton-X

added

Fig. 6. Cell reactions captured with our impedance-to-frequency converter(black solid line, left axis) and actual impedance at 10 kHz measuredsimultaneously via a LCR meter connected to on-chip electrodes (red solidline, right axis). Also shown are data from a control experiment without cells(dashed lines).

Frequency [Hz] Frequency [Hz]

Fig. 7. Results of impedance spectroscopy with and without cells obtainedduring the experiment shown in Fig. 6, with added fitting curves

base level of the control experiment followed as the cells werekilled and detached from the chip.

Over the whole course of the experiment, the curves of thechip frequencies and the impedance measured by the LCRmeter show remarkable agreement, demonstrating the viabilityof our impedance-to-frequency converter for cell adhesionmeasurement. The slight discrepancies can be sufficientlyexplained by the facts that the two curves represent differentphysical measurement locations on the chip and that thefrequency has no one-to-one mathematical relation to theimpedance magnitude.

Fig. 7 shows data from the impedance spectroscopy before(i.e. with cells, denoted with squares) and after addition ofTriton-X (without cells, denoted with asterisks). The dashedlines represent our best fit of the model (1) for the electrodewithout cells, with the addition of a real constant term R toaccount for the spreading resistance through the medium. Ascan be seen, the model fits the measured data very well. Wealso performed a least squares fit of the impedance spectrum

Fig. 8. Chip photograph of cells growing on our chip roughly 20 hoursafter seeding. In the right half of the picture, a scratch in the cell monolayerexposes a row of electrodes.

with cells, represented by the solid lines. The real part ofthe impedance clearly shows the increased constant resistivepart predicted by our finite elements modeling. Scaling it fora single electrode yields a value of about 110 kΩ, in goodagreement with the expected value. Above 10 kHz, capacitivecoupling through the cells results in a drop of the in-phaseimpedance. At the same frequency, the imaginary part shows adeviation from (1) which we also approximated with a constantterm in our least squares fit. This behavior is assumed to resultfrom the constituent parts of the cell, whose effects are notaccounted for in our models.

B. Localized detection of adhesion

To demonstrate the capability of our chip to deliver aspatially resolved measurement of cell adhesion, we seededcells and again let them grow for about 20 hours until aconfluent monolayer had formed and the impedance signalshad stabilized. Then, we briefly removed the chip from theincubator and scratched the cell monolayer using a plasticpipette tip (Fig. 8). Fig. 9 shows the mean frequencies ofthe undisturbed electrodes and of the electrodes which wereexposed to the electrolyte, both normalized to the last datapoint before the scratch.

While the signal for the electrodes which were still coveredby cells shows no discernible change after the interruption,the frequencies of the exposed electrodes drop sharply as ex-pected. During the following hours, the frequencies graduallyrise again as cells along the edge of the scratch spread out,however the signal does not fully recover to the previous value.

VI. CONCLUSION

Based on a model of the electrode-electrolyte interface andthe impedance generated by a cell monolayer, we developed anovel, fully CMOS integrated biosensor for detecting cellularadhesion. Experimental results demonstrate the capability ofthe sensor to detect gradual cell responses to stimuli as well aslocalized variations in cell adhesion. Its measurement principlewith digital frequency-modulated output makes it easy to

exposed electrodes

Fig. 9. Normalized chip frequencies of an experiment where a confluent cellmonolayer was deliberately damaged at t = 0 (see Fig. 8). The dotted linerepresents the undisturbed electrodes, the solid line represents the electrodesthat were exposed.

combine with on-chip digital signal processing, as only acounter is needed.

In the final multi-parametric sensor chip, the capability todetect cell coverage locally will greatly enhance the informa-tion obtained from the other sensors, e.g. for bioluminescenceand cell metabolic activity. This is expected to enable thedetection of a wide range of possible contaminants withunprecedented reliability.

ACKNOWLEDGMENT

We sincerely thank Ulrich Bohrn for his tireless support inconducting the biological experiments.

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