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Sensors and Actuators B 199 (2014) 470–478

Contents lists available at ScienceDirect

Sensors and Actuators B: Chemical

jo u r nal homep age: www.elsev ier .com/ locate /snb

ll-plastic, low-power, disposable, continuous-flow PCR chip withntegrated microheaters for rapid DNA amplification

espina Moschoua, Nikolaos Vourdasa, George Kokkorisa, George Papadakisa,ohn Partheniosb, Stavros Chatzandroulisa, Angeliki Tserepia,∗

Institute of Nanoscience & Nanotechnology, NCSR “Demokritos”, P.O. Box 60037, 153 10 Ag. Paraskevi, GreeceInstitute of Chemical Engineering Sciences (ICE-HT), FORTH, Stadiou Str., Platani, P.O. Box 1414, 26504 Patras, Greece

r t i c l e i n f o

rticle history:eceived 12 December 2013eceived in revised form 1 April 2014ccepted 3 April 2014vailable online 15 April 2014

eywords:icrofluidics

ontinuous-flowicroPCR

olyimidelexible substrates

a b s t r a c t

The design, fabrication and evaluation of a low-cost and low-power, continuous-flow microfluidic devicefor DNA amplification by polymerase chain reaction (PCR) with integrated heating elements, on a com-mercially available thin polymeric substrate (Pyralux® Polyimide), is presented. The small thermal massof the chip, in combination with the low thermal diffusivity of the polymeric substrate on which theheating elements reside, yields a low power consumption PCR chip with fast amplification rates. A flow-through �PCR device is designed and fabricated using flexible printed circuit (FPC) technology on afoot-print area of 8 cm × 6 cm with a meandering microchannel realized at a very small distance (50 �m)above 3 independently operating resistive (copper) serpentine microheaters, each one defining one ofthe three PCR temperature zones. The 145 cm-long microchannel is appropriately designed to cross thealternating temperature zones as many times as necessary for the DNA sample to perform 30 PCR cycles.Numerical computations lead the design so that there is no thermal crosstalk between the 3 zones of ourchip and indicate excellent temperature uniformity in each zone. In addition, the total power consump-tion during the chip operation is calculated to be in the order of a few Watts, verified experimentally

by means of thermal characterization of our heaters. Thermal camera measurements also verified theexcellent temperature uniformity in the three thermal zones. An external, home-made temperature con-trol system was utilized to maintain the heater temperatures in the designated values (±0.2 ◦C). The PCRchip was validated by a successful amplification of a 90 base-pairs DNA template of the mouse GAPDHhousekeeping gene within 5 min.

© 2014 Elsevier B.V. All rights reserved.

. Introduction

Micro-total analysis systems has been a research field of increas-ng interest in the past years, enabling biochemical analysis foroint-of-care (POC) applications, as it favorably combines advan-ages such as low reagent consumption and short analysis time,n devices of small footprint, and thus of increased portability andow fabrication cost [1,2]. Miniaturized polymerase chain reaction�PCR) devices are expected to become a central part of most lab-

n-a-chip (LOC) or POC systems intended for molecular or clinicaliagnostics [3], and food safety monitoring [4,5]. PCR, a widespreadiochemical process for amplifying trace amounts of DNA in a

∗ Corresponding author. Tel.: +30 210 650 3264.E-mail address: atserepi@imel.demokritos.gr (A. Tserepi).

ttp://dx.doi.org/10.1016/j.snb.2014.04.007925-4005/© 2014 Elsevier B.V. All rights reserved.

sample to generate copies sufficient for detection, is based onrepeated thermal cycling of the DNA sample through three tem-perature steps (denaturation, annealing, and extension). After eachcycle, the number of the DNA copies can be doubled, and thus20–30 cycles can lead to millions of DNA copies. PCRs performed inmicrofluidic devices provide much faster processing, in the orderof minutes rather than hours, compared to conventional thermocy-clers [6], while allowing the production of highly integrated devices[7].

Two basic types of designs have been followed by researchers forthe realization of �PCR devices: stationary chambers with cyclingtemperature, or continuous-flow devices typically with three zonesmaintained at constant temperatures and only the sample chang-

ing temperature as it flows through the zones. �PCRs with staticchambers were historically the first to be realized [8], howevercontinuous-flow devices have been proven faster [9] than the staticones. The choice of the type of �PCR design drives in many cases the

d Actuators B 199 (2014) 470–478 471

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D. Moschou et al. / Sensors an

hoice of the material to be used for the fabrication of �PCR devices,n order to best exploit the advantages of each material for improv-ng the device operation. For example, Silicon (Si) [8] has preferablyeen used for static devices, where the high thermal conductiv-

ty of the material favors fast heating or cooling rates, however toaintain high temperature uniformity within the sample, thermal

solation of the device is necessary. On the other hand, materi-ls with low thermal conductivity, such as glass [9] or plastics10], are preferable for continuous-flow devices, where tempera-ure cycling of the entire device (large thermal mass) is avoided,hile the temperature uniformity and low power consumption

s ensured, under conditions, by the good thermal isolation pro-ided by these materials. Plastics, such as polydimethylsiloxanePDMS) [11], polymethylmethacrylate (PMMA) [12], polyimide (PI)13–15], and printed-circuit-boards (PCB) [16] have a cost advan-age over Si and glass, as well as ease of processing [17]. The lowerabrication cost of plastic �PCR devices combined with the capabil-ty for mass production of such devices makes plastics the materialf choice for the fabrication and commercialization of �PCRs as wells their capability to be integrated into larger more complicatedOC systems [3,7,18].

Another important feature of �PCRs is the heating systemequired for performing PCR. Although external systems such asR lamps [13], heating blocks or Peltier elements [11,19] have been

ostly used in �PCRs, integration of microheaters and tempera-ure sensors on the device becomes increasingly popular, due tohe advantage it offers for a more compact and thus of increasedortability system. Thin film microheaters are usually depositedn glass [20], and thus most of the �PCRs integrated with micro-eaters are hybrid devices [10,21,22] combining Si or plastic andlass, therefore hindering the mass production of such devices.

In this work, an all-plastic continuous-flow �PCR devices designed and fabricated with integrated heating elements,n a commercially available, biocompatible [23,24], very thin100 �m) PI substrate (Pyralux®), utilizing FCB-compatible tech-ology, amenable to mass production. To the best of our knowledge,olyimides have been used so far as substrates either for the real-

zation of static �PCR chambers [13], or for the fabrication ofesistive heating elements bonded to glass-based static chambers14]. In spite of the undisputed advantages of continuous-flowPCR devices, such devices have not yet been demonstrated onI, and more specifically on Copper(Cu)-clad PI that allows thentegration of microfluidic channels and resistive heating ele-

ents on the same substrate [25], without requiring an additionaletal deposition step. Furthermore, the use of a very thin, ther-ally isolated PI substrate is expected to promote further the

apability for even lower power consumption required for theevice thermal operation, increasing the possibility of making aattery-operated and portable �PCR. Numerical calculations areerformed to (a) demonstrate the advantages of a thin plastic sub-trate for �PCR, in terms of power consumption, thermal isolationetween the PCR zones and temperature uniformity of the zonesnd (b) suggest suitable respective operating parameters. Finally,he fabricated �PCR is validated for successful and rapid DNAmplification.

. Materials and methods

.1. Chip design

A schematic of the continuous flow �PCR device comprising

oth a microfluidic circuit and resistive microheaters is illus-rated in Fig. 1. This device, which is an improved version of aesign presented previously [25], implements 30 thermal cycles

n total for efficient amplification of DNA. Each cycle features three

Fig. 1. A 3d schematic of the proposed flexible �PCR chip comprising a meanderingmicrochannel (shown in blue) and resistive microheaters (shown in brown). Onlythree unit cells, each one implementing a thermal cycle, are shown.

thermal steps, hence three discrete thermal zones are designed,each controlled by an individual resistive microheater.

The meandering microchannel, of a total length of 1.45 m,where the DNA sample flows continuously, is designed to crossthe three alternating PCR temperature zones, so that the arclengthalong the channel axis, l, lying above each one of the zones isthe same (lde = lan = lex = 1.2 cm). Since the device is designed toaccommodate PCR protocols with relative time ratios of 1:1:2 fordenaturation:annealing:extension, the channel width at the exten-sion zone is designed twice as much of that at the denaturation andannealing zones, so that the DNA sample residence time in the threezones follows the same ratio, i.e., 1:1:2. Thus, a width of 400 �m waschosen for the extension zone microchannel, and 200 �m for thedenaturation and annealing microchannels. The choice of doublingthe extension microchannel width as well as using a microchannelminimum feature size at 200 �m further optimized our previousdesign [25], facilitating the chip fabrication and achieving lowerpressure drop across the channel, thus increasing the device sealingdurability.

The integrated resistive microheaters are designed to be fabri-cated at the bottom Cu layer of the commercially available Cu-cladPI substrate (Fig. 1). The resistors are also meandering, so as toprovide the largest possible electrical resistance in the space avail-able for each heating zone, thus ensuring sensitive temperaturecontrol. The Cu lines are 100 �m in width, 2 m in length for denatur-ation, 2.1 m for extension, and 1.5 m for the annealing microheater.The separation between the heating zones is determined throughsimulation for minimal thermal cross-talk. According to the geo-metrical characteristics of the Cu lines, their electrical resistance iscalculated to be 23 � for denaturation/extension, and 30 � for theannealing microheater.

2.2. Heat transfer and fluid flow calculations in the device

The design of the �PCR device was assisted by simulations aim-ing at calculating (a) the temperature uniformity in each one ofthe �PCR zones, (b) the pressure drop in the microfluidic channel,and (c) the power requirements for heating the device. Regardingthe first aim of the simulation, the requirement for a high per-formance �PCR device is uniform temperature of the fluid (DNAsample) within the zones: The fluid temperature variation shouldbe less than 3 K [26] (±1.5 K) in the denaturation, the annealing,and the extension zone. Given that the fluid conveys heat from onezone to the other, the simulation suggests the suitable operating

conditions to inhibit thermal “cross-talk” between the zones. Theadvantage of using polymeric layers instead of SiO2 or Si layers asthe device substrate was demonstrated in a previous work [27];the temperature uniformity was much better in the device with

472 D. Moschou et al. / Sensors and Actuators B 199 (2014) 470–478

abrica

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npc

ihs

∇a

∇w(flb(0

�

∇wiPacc(padtf

Fig. 2. �PCR chip f

he polymeric substrate. The second aim is the calculation of theressure drop in the microfluidic channel which is critical to pre-ent delamination of the cover layer sealing it. The lower theressure drop, the lower the bonding strength required to keep theevice leak-tight. The third aim of the simulation is the calculationf the power requirements for heating the device.

During simulations, heat transfer and fluid flow equations areumerically solved in the unit cell of the device (Fig. ESI-1a, sup-lementary material) where one thermal cycle occurs. The deviceonsists of 30 cells (equal to the number of cycles) in an array.

Concerning the mathematical formulation, the temperature, T,s calculated at steady state by solving in 3d the equations for theeat transfer in the polymeric layers of the unit cell (Fig. ESI-1b,upplementary material)

· (−kp∇T) = 0, (1)

nd for the heat transfer in the fluid,

· (−kf ∇T) = −�Cpu · ∇T (2)

here kp and kf are the thermal conductivities of PI and the fluidDNA sample). � and Cp are the density and the specific heat of theuid. u is the fluid velocity which is calculated by the momentumalance and continuity equations [28] for steady state laminar flowthe Reynolds number for the conditions of the simulations is from.11 to 5, see supplementary material),

u · ∇u = ∇ ·{

−pI + �[∇u + (∇u)T ] − 23

�(∇ · u)I}

+ �g (3)

· (u) = 0 (4)

here � is the dynamic viscosity, p is the pressure of the fluid, Is the unit (identity) matrix and g is the gravitational acceleration.eriodic boundary conditions are considered for Eqs. (1) and (2)t the boundaries of the unit cell which are in contact with adja-ent cells. At all other boundaries of the unit cell, heat transfer byonvection to the ambient is considered. Concerning Eqs. (3) and4), no slip condition is considered at the channel walls. The tem-erature in the heaters is considered uniform and is calculated by

trial and error procedure so as the average temperatures at theenaturation, extension, and annealing zones are equal to the per-inent set points (368 K, 345 K, and 328 K). The power requirementsor the heating of the device are the power losses to the ambient.

tion process flow.

The numerical solution is performed by the finite element methodimplemented by the commercial code COMSOL [29].

The thermal conductivity of PI is set at 0.12 W/(m K) [30]. Thefluid properties are those of water; the density, dynamic viscos-ity, thermal conductivity, and specific heat are coming from thedatabase of the commercial code and are functions of temperature.The latter means that Eqs. (1)–(4) are coupled. For the base case cal-culations, the heat transfer coefficient to the ambient is 5 W/(m2 K)and the ambient temperature is 293 K (20 ◦C).

2.3. Fabrication of the �PCR device

The process flow followed for the fabrication of the all-polymer�PCR chip is schematically depicted in Fig. 2. It begins with a com-mercially available Cu-clad substrate and by means of standardphotolithography to pattern the microfluidic on the top Cu layer.The chosen substrate for our flow-through �PCR chip is a com-mercially available flexible PI film (Pyralux®, Laminate AP 8545R,sheets with 100 �m PI) with 18 �m rolled-annealed Cu foil on bothsides of PI (purchased from DuPontTM), allowing the direct inte-gration of resistive microheaters on the same substrate with themicrofluidics, without any further processing steps for metal depo-sition or attachment of external heating elements. By means oftypical Cu wet etching (in commercial FeCl3 solution), the chan-nel is irreversibly formed on Cu (Fig. 2 step 1). Next, standardphotolithography and Cu wet etching follow for the formation ofmicroheaters on the bottom Cu layer (Fig. 2 step 2). Utilizing thepatterned top Cu layer as an etch mask, the meandering microchan-nel is plasma etched in the underlying PI layer (Fig. 2 step 3),reaching a depth of 50 �m. Plasma etching of PI was performed withO2/SF6 processing mixture, in a high density plasma etcher (AlcatelMET), for the formation of the microfluidic channels. The heliconsource (at 13.56 MHz) provides RF power (up to 2000 W) to anAg-plated Cu antenna. The patterned PI substrates were thermally‘glued’ to Si wafers with conductive paste to ensure the heat transferfrom the processed substrate to the He-backside cooled Si carrier,thus maintaining the Pyralux® temperature at the desired level. By

means of a second capacitively coupled RF generator, the samplewas independently polarized to control the ion-bombardment onthe surface. The plasma process parameters were optimized [25]for maximum PI etching rate.

D. Moschou et al. / Sensors and Actuators B 199 (2014) 470–478 473

Fig. 3. Photograph of a fabricated �PCR device. The three temperature zones,defined by the three Cu microresistors beneath the meandering microchannel, arestr

eswist

ipft

2

hSbtpt

�Twottattmim

2

fgpfsme(

Fig. 4. Temperature at xy cross sections of the unit cell at the middle of the microflu-idic channel height; the focus is on the region between heater 1 (denaturation) andheater 2 (extension) [see (a) and (b)] and heater 2 and heater 3 (annealing) [see

hown. In the inset, a part of the device is shown in magnification, where details ofhe microchannel and the Cu microheaters are visible at the top and bottom side,espectively.

After formation of the microchannel in PI, the top Cu layer istched away (Fig. 2 step 4) and the microchannel is sealed (Fig. 2tep 5) by laminating a 70 �m thick commercial Kapton® PI tapeith a silicon adhesive. Prior to the lamination, a mild reactive

on etching (RIE) plasma treatment is performed on both the sub-trate and the cover layer (100 mTorr, 50 sccm O2, 100 W, for 30 s)o enhance the bonding strength and provide irreversible sealing.

The final chip (Fig. 3) size is about 8 cm × 6 cm. Sealed microflu-dic channels were tested for leakage using a laboratory syringeump (Chemyx Inc., Fusion 200). The fluidic interfacing was per-ormed with commercial Upchurch nanoport fittings adhered onhe chip.

.4. Experimental set-up for microheater characterization

In order to characterize and calibrate the fabricated micro-eaters, a probe station was used with a heated base (Signatone-1041-D3 Hot chuck) and a Keithley 2400 source-meter controlledy Labview, for sourcing current I and measuring electrical resis-ance R. The heaters were glued on a wafer substrate using thermalaste, which was in-turn vacuum attached to a hot plate of con-rolled temperature.

For independent temperature measurements on fabricatedPCRs during microheater operation, a thermal camera (NIKONhermal Vision LAIRD 3A) was utilized. First, the emissivity of the PIas calculated. The PI substrate was partially coated with a material

f known emissivity (black Krylon paint, emissivity = 0.97). Then,he PI was attached on a hot plate with thermal paste to takehermal camera images for different temperatures. Knowing thebsolute temperature and the emissivity of the material (Krylon),he PI emissivity was calculated to match the temperature value,hrough the thermal camera image analysis software. The deter-

ined PI emissivity was used for the calibration of the thermalmages/temperatures obtained during operation of the integrated

icroheaters.

.5. PCR sample preparation and evaluation

Evaluation of our device for DNA amplification was per-ormed with a 92-bp DNA fragment from the mouse housekeepinglyceraldehyde 3-phosphate dehydrogenase (GAPDH) gene. Therimers that were used to produce this fragment were purchasedrom Microchemistry (Crete, Greece) comprising the following

equences: mGAPDH F: 5′-ACC ACA GTC CAT GCC ATC AC-3′ andGAPDH R: 5′-TCC ACC ACC CTG TTG CTG TA-3′. 10 pmoles of

ach primer were mixed with KAPA 2G Fast DNA polymeraseKAPABIOSYSTEMS) according to the manufacturer protocol. Each

(c) and (d)] for two volumetric flow rates equal to 1.8 �l/min [see (a) and (c)] and7.8 �l/min [see (b) and (d)]. The direction of fluid flow is indicated with arrows.

PCR reaction was supplemented with MgCl2 and BSA to a finalconcentration of 2.5 mM and 100 �g/ml, respectively. To avoidadsorption of biological sample on microchannel walls duringactual �PCR experiments, the blocking agent BSA was added tothe �PCR mixture. The amplification protocol used with the con-ventional thermocycler consisted of 1 min initial denaturation at95 ◦C, 25 cycles of 10′′ denaturation at 95 ◦C – 10′′ annealing at 55 ◦C– 1′′ extension at 72 ◦C and a final 1 min extension at 72 ◦C. PCRproducts were loaded on a 2% agarose gel stained with ethidiumbromide and visualized with an ultraviolet (UV) transilluminator.The 2% agarose gel was prepared by mixing 2 g of agarose (SeaKemLE agarose, CAMBREX) in 100 ml of 0.5× TBE (10× TBE: 890 mMTris, 890 mM Boric acid, 20 mM EDTA, pH at 25 ◦C: 8.0). The solu-tion was melted for 3 min into a microwave and let it cool down.10 �l of EtBr (10,000×) were added and the gel was poured into theelectrophoresis tray. 25 �l of the collected flow-through amplifiedDNA were mixed with 6× orange loading dye and loaded on thegel. Electrophoresis was performed using a Biorad power supplysystem for 1 h at 100 V. 5 �l of Puc19 vector digested with BsiS Irestriction enzyme (50 ng/�l, Minotech, Heraklion, Crete) was usedas a reference DNA marker.

3. Results and discussion

3.1. Simulation results

The temperature at xy cross sections of the unit cell at the middleof the height of the channel is shown in Fig. 4. The cross sections inFig. 4a and b shows the temperature in the region between heater 1(denaturation) and heater 2, for two volumetric flow rates, and fora separation of 1.1 mm between heater 1 and 2. The cross sectionsin Fig. 4c and d focus on the region between heater 2 (extension)and heater 3 (annealing), for a separation of 1.7 mm. The greaterthe volumetric flow rate is, the greater is its effect on the tem-perature distribution. When the volumetric flow rate is 1.8 �l/min(Fig. 4a and c), although the fluid flow affects the temperature at the

regions between the heaters, there is no significant thermal “cross-talk” between the zones. Thermal “cross-talk” appears when thevolumetric flow rate is 7.8 �l/min (Fig. 4b and d).

4 d Actuators B 199 (2014) 470–478

aopTIaislii“baat

flicoealcTri

trrd(oaF[Tac

3

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R

wtcfR0

Fig. 5. (a) The percentage of the volume of each zone in the acceptable temperaturerange (set point ±1.5 K) vs. the volumetric flow rate. (b) The pressure drop in a unit

74 D. Moschou et al. / Sensors an

In order to quantify the effect of thermal “cross-talk” on temper-ture uniformity within each zone, the percentage of the volumef a zone in the acceptable temperature range (±1.5 K from the setoint) vs. the volumetric flow rate is calculated and shown in Fig. 5a.he percentage decreases almost linearly with flow rate in all zones.n the denaturation zone, there is a steep decrease of the percent-ge value at around 8 �l/min. The reason for this steep decreases that a large fraction of the volume in the denaturation zonelightly exceeds the temperature upper limit (set point +1.5 K); thisarge fraction exceeds 369.5 K by less than 0.3 K. The reason for thencrease of the fluid temperature, at higher volumetric flow rates,s that when the fluid velocity in the channel increases, thermalcross-talk” increases and the temperature at the heaters shoulde also increased so as to get the average temperature of the fluidt the set point. Fig. 5a shows that it is beneficial for the temper-ture uniformity at each zone to use a volumetric flow rate lowerhan 8 �l/min.

In Fig. 5b, the pressure drop in the unit cell vs. the volumetricow rate is shown. In order to calculate the total pressure drop

n the device, the values should be multiplied by the number ofycles (30 in this case). For the range of volumetric flow ratesf Fig. 5b, which is wider than the range finally used in the PCRxperiment, the maximum pressure drop across the �PCR chip ispproximately 60 kPa. The calculated maximum pressure drop isower than typical bonding strengths achieved with irreversible,hemically-assisted sealing processes of polymeric substrates [31].herefore, even high volumetric flow rates are not expected toisk the integrity of the designed �PCR device, provided thatrreversible sealing processes are implemented.

The power required for maintaining the unit cell at the desiredemperature set points is not affected significantly by the volumet-ic flow rate: It is calculated to vary from 41.5 to 42.2 mW for flowates varying from 0.6 to 11.4 �l/min. However, the required powerepends strongly on the heat transfer coefficient at the ambienth) and the ambient temperature (Tamb). The power requirementsf the unit cell during device operation for different values of thembient temperature and heat transfer coefficient are shown inig. 5c: the calculated power for the whole device is from 1.14h = 5 W/(m2 K) and Tamb = 298 K] to 2.66 W [h = 10 W/(m2 K) andamb = 288 K]. The value determined by the experiment, which isround 2.4 W (see Section 3.2), is in the range predicted by thealculations.

.2. Microheater characterization

After the fabrication of the �PCR chip and its testing under liq-id flow, the calibration of the integrated resistors to be used asemperature sensors was the next step. To this end, the fabricatedesistors were placed on a hot plate and their electrical resis-ance for varying temperatures within the range of interest wereecorded. In Fig. ESI-2 of supplementary material, a representativealibration curve of measured resistance R versus the temperature

is shown, on one of the fabricated chips. The measurements wereepeated for several other chips and average values were deter-ined. The linear behavior of R with T, for the PCR desired range of

emperatures, allows for the utilization of the resistors as temper-ture sensors, according to [14,32]:

= R0 + R0 · TCR(T − T0)

here R is the electrical resistance at a temperature T, R0 the resis-ance value at a known temperature T0 and TCR the temperature

oefficient of resistance. The TCR is a material related property andor Cu is ∼0.0039 ◦C−1 [33]. By linear fitting of the experimental–T curve with the aforementioned equation, an average TCR of.0032 ◦C−1 ± 0.0005 is estimated for our fabricated resistors, to be

cell vs. the volumetric flow rate. For the calculation of the total pressure drop inthe device, the values should be multiplied by 30 (number of cycles). (c) The powerrequirements for heating the unit cell during device operation for different values ofambient temperature and heat transfer coefficient. Volumetric flow rate is 3 �l/min.

D. Moschou et al. / Sensors and Actuators B 199 (2014) 470–478 475

0.4 0.6 0.8 1.0 1.2 1.4 1.6 1.850

60

70

80

90

100

110

I1=225 mAI2=185 mAI3=164 mA

T3

T2

T (o C

)

P (W)

T1

Fig. 6. Microheater temperature T as a function of the supplied power P for a rep-rs

uoso

odmiwfeezwprsccttaimt(d

miPhmwuTstcfti(wa

Fig. 7. (a) Thermal camera image of a chip, showing the meandering Cu micro-heaters, operating at typical PCR temperatures. Depicted are the extension and

achieved by controlling the voltage across the resistive heaters

esentative fabricated Cu microheater. Respective injected current values are alsohown to achieve the required PCR temperatures (T1, T2, T3).

sed for determining the temperature by measuring the resistancebtained by the temperature control system. The average value andtandard deviation of the TCR were obtained from measurementsn 6 separately fabricated integrated microresistors.

After the calibration of the fabricated microheaters, evaluationf their performance as heating elements followed. This involvedriving and monitoring current through each microheater whilsteasuring the corresponding voltage drop across it to calculate

ts resistance at steady state. For this reason, electrical interfacingas made through a flexible cable connector fitted in the Cu pads

ormed on the flexible chip, connecting the microresistors to anxternal current source system, providing the required power forach individual resistor to heat the above-lying PCR temperatureone. By utilizing the previously estimated TCR, their resistanceas translated to temperature and plotted against the electricalower required to achieve these temperatures. In Fig. 6, the powerequired by a fabricated resistor to reach the PCR temperatures ishown. In the same graph, the supplied currents and thus the cal-ulated power consumption are shown, from which a total poweronsumption of 2.7 W is estimated. For simultaneous operation ofhe three microresistors, the total power consumption is reducedo 2.4 W [25]. Compared to other devices presented in the liter-ture [14,33], the present device requires lower power, due tots thin substrate material, and thus its low thermal mass. The

easured power consumption is in agreement with the simula-ion results (see Section 3.1), given the fact that some power lossparasitic resistance) is expected in the electrical interfacing of theevice.

As the absolute values of the temperature achieved on theicroresistors are crucial for successful PCR reactions, they were

ndependently measured directly on the device, and especially onI, which is the material in contact with the DNA sample to beeated. For this reason, the temperature of the three zones undericroresistor operation was measured, using the thermal camera,hile at the same time recording the corresponding resistor val-es and calculating the temperature by taking into account theCR of Cu. First, the emissivity of PI (see Section 2.3) was mea-ured and found to be approximately 0.97, so as to calibrate thehermal camera readings. A representative thermal camera imagean be seen in Fig. 7a, showing a part of our heated �PCR chiprom the side of the Cu microheaters (without fluid passing throughhe chip) between the extension and annealing zones. The imagendicates that the temperature is very uniform across the zones

after statistical analysis across each zone, a variation of ±0.2 ◦Cas determined), while no noticeable thermal crosstalk between

djacent zones was observed. The different colors observed on

annealing zones. (b) Absolute temperature measured (Tmeasured) by a thermal cam-era on PI surface around the microheaters versus the microheater temperature(Testimated) as estimated from its resistance value and Cu TCR.

Cu heaters and the PI substrate are attributed to the differentemissivity of the two materials (about 0.6 for Cu and 0.97 for PI).In Fig. 7b, the temperatures achieved on the microresistors areshown, as measured by the thermal camera, versus temperatureas estimated using electrical resistance and Cu TCR. Indeed, forall resistors, a linear relationship was observed between measured(thermal image) and estimated temperatures, with very small dif-ferences (less than 3 ◦C) between their values (more prominent atlower temperatures).

3.3. Temperature control

In order to ensure temperature stability of the microheatersat temperatures necessitated by the PCR protocol, a temperature-control unit was required. The microheaters have a dual role andare used both as heating elements as well as temperature sensors toprovide feedback for the temperature control system, associatingtheir electrical resistance value to their temperature.

To control and monitor the �PCR operation, a specially designedtemperature controller board was developed. It is built around amicro-controller which is responsible for controlling and readingthe onboard DACs and ADCs and communicate with a personalcomputer (PC). The circuit is able to control up to three integratedCu micro-heaters necessary for the �PCR. Temperature control is

whilst a small sensing resistor is used to measure the current flow-ing through them. In order to increase the accuracy of the controlloop the driving of the micro-heaters is performed via an external

4 d Actuators B 199 (2014) 470–478

1nactm

tosttotsraacPa((

3

piaswflartp

avdrdssoiir1t72pouoaaeprogP

Fig. 8. (a) Experimental setup for testing the �PCR chip for DNA amplification. (b)Image of the gel after electrophoresis of three PCR samples. The first and secondlanes correspond to �PCR DNA products produced under flow speed of 5 �l/min

76 D. Moschou et al. / Sensors an

6-bit DAC and a dedicated power amplifier in order to supply theecessary current. For the same reason an external 16-bit ADC islso used for reading the sensing resistor. The controller board isonnected via the USB port with a PC in which the software forhe PID (proportional-integral-derivative) control loop is imple-

ented.Following the calibration of the resistors for use as tempera-

ure sensors and the measurement of the power required for eachne to achieve the desired temperatures, the temperature controlystem was tested, in respect to its ability to achieve and main-ain stable temperature in the three heating zones, as required forhe PCR protocol. The temperature of a heater vs. time during theperation of the controller is shown in Fig. ESI-3 of supplemen-ary material. In this graph, the controller changes the set point atpecific time instances. It is observed that the controller can accu-ately set the temperature of each zone at the specified value. Inddition, the lack of fluctuations over time (in the order of ±0.2 ◦Ct steady state) and the fast response of the heaters to temperaturehanges insures reliable operation and thus can be used reliably forCR experiments. Furthermore, very fast heating and cooling ratesre obtained during temperature transitions from 55 ◦C to 72 ◦C19.5 ◦C s−1) and 72 ◦C to 95 ◦C (12.5 ◦C s−1), and from 95 ◦C to 55 ◦C7.2 ◦C s−1).

.4. On-chip DNA amplification

To conduct the PCR experiment on-chip, PCR mixtures wererepared in order to produce a 92-bp DNA fragment correspond-

ng to the mouse GAPDH housekeeping gene. BSA (10 mg/ml) wasdded to the mixture in a final dilution of 1:100 for channel pas-ivation and enhancement of the PCR reaction. The PCR mixtureas injected in the chip utilizing a peristaltic pump (Fig. 8a), underow rates ranging from 1 to 5 �l/min. Between different PCR runs,

buffer solution was passed through the device to assure that noesidual DNA contaminated the next sample. Also, a negative con-rol with no template was passed, to assure our results were not aroduct of residual contamination.

�PCR products were collected at the outlet and analyzed bygarose gel electrophoresis, comparing them to the respective con-entional thermocycler results. In Fig. 8b, the agarose gel imageisplaying the DNA amount produced under two different flowates with the �PCR device and its comparison with the amounterived from the thermocycler is shown. DNA amplification wasuccessfully achieved in our chips for both flow rates, utilizing 25 �lamples and delivering their product in 5 and 10 min, for flow ratesf 5 �l/min and 2.5 �l/min, respectively. As indicated by the gelmage, faster flow rates yield lower PCR product. Indeed, a 15%ncrease in DNA product was observed with 2× decrease of flowate. The observed DNA amplifications were achieved in approx.0 s and 20 s per cycle, for a flow of 5 �l/min and 2.5 �l/min, respec-ively. The estimated amount of produced DNA is approximately5 ng with a flow rate of 5 �l/min, and 15% more with a flow rate of.5 �l/min. The conventional PCR resulted in at least 4 times moreroduct. The amount of the starting template used was in the rangef 1–10 ng. The signal to noise ratio observed for the DNA prod-ct obtained with a flow rate of 5 �l/min indicates the potential ofur device to satisfactorily amplify smaller sample volumes, thuschieving faster detectable PCR products. The shortest time for DNAmplification in �PCR reported until now, to the best of our knowl-dge, is 4.5 s/cycle, for a �PCR on a glass substrate [9], while recentlyresented polymer �PCR devices with higher power consumption

eport minimum cycle duration of about 6 or 10 s/cycle [34,35]. Anptimized protocol for the �PCR chip is currently under investi-ation in order to achieve yields comparable to the conventionalCR.

and 2.5 �l/min, respectively. On the third lane, 1/2 of the respective PCR productderived from the thermocycler is shown. DNA marker Puc19/BsiS I is shown in lane4.

4. Conclusions

The design and fabrication of a continuous-flow all-polymer�PCR chip is presented. Our chip combines the advantages of fastamplification rates, small thermal mass and therefore low powerconsumption (approximately 2.4 W), as well as the potential ofbeing integrated in more complex LoC systems. The chip designwas assisted by heat transfer and fluid flow calculations to ensuretemperature uniformity across the PCR temperature zones andminimal cross-talk between the zones, for a range of recommendedoperational parameters such as the volumetric flow rate. The chipis fabricated in a biocompatible, commercially available thin PIsubstrate with integrated Cu microheaters on its bottom side.The microheaters are successfully implemented both as temper-ature sensors and as heating elements, controlled by a home-madetemperature-control system. The temperature uniformity and thepower consumption of our chip were verified experimentally, ingood agreement with the simulation results. Finally, successfulDNA amplification was demonstrated within 5 min, placing ourchip among the fastest ever reported [36].

Acknowledgements

This work was co-financed by Hellenic Funds and by the Euro-pean Regional Development Fund under the NSRF 2007–2013,

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ccording to Contract no. MICRO2-45 of the Project “Microelec-ronic Components for Lab-on-chip molecular analysis instrumentsor genetic and environmental applications”. Two of the authorsG.P. and G.K.) acknowledge support of the DoW (LS7-276) ProjectSupporting post-doctoral researchers”, Ministry of Education, Life-ong Learning, and Religious Affairs); the source of funding ishe European Social Fund (ESF) – European Union and Nationalesources.

ppendix A. Supplementary data

Supplementary data associated with this article can be found, inhe online version, at http://dx.doi.org/10.1016/j.snb.2014.04.007.

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Biographies

Dr. Despina Moschou is an Electrical and Computer Engineer, graduate of theNational Technical University of Athens (2005). She received her Ph.D. from theNational and Kapodistrian University of Athens, Department of Informatics andTelecommunications (2009). During her Ph.D., she conducted research at theInstitute of Microelectronics (IMEL), NCSR “Demokritos”, working on the fabri-cation, characterization and reliability of novel thin film transistors (TFTs). Since2010, she is a post-doctoral collaborator at IMEL, NCSR “Demokritos”, working onintegradable polymer microfluidic chips and Lab-On-a-Chip systems for bioanalyt-ical applications. She is the author/co-author of 37 papers presented in internationalpeer-reviewed journals and conferences.

Dr. Nikolaos Vourdas holds a Diploma in Chemical Engineering (2001) and a M.Sc.in Materials Science and Technology (2003), both from National Technical Univer-sity of Athens (NTUA). He completed his Ph.D. Thesis (2007) at the National Centerfor Scientific Research (NCSR) “Demokritos” on plasma processing of polymers formicro-fabrication and surface modification. His Ph.D. thesis won the “Best DoctoralThesis Award of the year 2007” at NTUA. After the end of his doctoral studies, heworked for the Hellenic Aerospace Industry (HAI). Currently, he is with the Instituteof Nanoscience & Nanotechnology (former IMEL). Dr. N. Vourdas is an author/co-author of 24 publications in peer reviewed journals and also holds two internationalpatents.

Dr. George Kokkoris holds a Diploma of Chemical Engineering from the NationalTechnical University of Athens (NTUA) (1998), a Master’s degree on Microelectron-ics (2000) from University of Athens, and a Ph.D. from NTUA (2005). His currentresearch is on the mechanisms of surface roughness formation during fabricationprocesses, on multiscale modeling of plasma etching and chemical vapor deposi-tion, as well as on modeling of microfluidic devices. He is working at the Institute ofNanoscience and Nanotechnology (former IMEL) of the National Center for ScientificResearch “Demokritos”. He is the author/co-author of more than 40 publications inpeer reviewed journals.

Dr. George Papadakis is a molecular biologist with a Ph.D. in acoustic biosensors.He has work experience related to molecular and biochemical analysis of biologicalsamples performed in diagnostic centers and research labs. He is currently involvedin the development of novel methods for microbial detection and screening for DNAmutations. He is also focused on the construction of lab-on-a-chip platforms usingacoustic wave sensors for diagnostic purposes. At present, he is a post-doctoral col-laborator at the Institute of Nanoscience & Nanotechnology (former IMEL), NCSR“Demokritos”, Athens. He is the author of 14 publications in peer-reviewed journalsand co-author of 1 granted patent.

Dr. John Parthenios is a Principal Scientist at FORTH/ICEHT since February 2009.He received his university degree in the Department of Physics of the University ofPatras in 1988, and his Ph.D. degree from the same department in Laser Physics. Hecarried out post- doctoral research for 9 years in FORTH/ICE-HT working mainly inthe field of smart materials and structures. During 2008, he was a visiting scientist

at University of London, Queen Mary and Westfield College. He has expertise ininstrumentation techniques for non-destructive testing of materials, targeting tothe use of spectroscopic technique such as Raman Spectroscopy for stress/strain andtemperature sensing at the micro and nano scale. Currently his research interests arein the thermomechanical characterization of nanomaterials (Graphene, GO, CNTs,

4 d Actu

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DAfNRloc

interests include microfluidic devices and wetting properties thereof, biomoleculemicroarrays, and LOC devices. She has served in the organizing committee of several

78 D. Moschou et al. / Sensors an

NT yarns and carbon fibers) and their nanocomposites. He was co-chairing therapHEL conference in 2013, is co-author of four chapters in books and has published2 papers in international journals (total number of citations: >800; h-index: 14)nd more than 100 presentations in national and international conferences. He haseen in the key personnel of over 15 research programs (funded by CEC, nationalodies and industry).

r. Stavros Chatzandroulis holds a B.Sc. from the Physics Department, University ofthens, Greece in 1990, a M.Sc. on Electronic Automation in 1993 and a Ph.D. degree

rom the same department in 1999. He is currently a researcher with the Institute of

anoscience & Nanotechnology (former IMEL), at the National Center for Scientificesearch “Demokritos”. His research interests focus on physical, chemical and bio-

ogical sensors and devices, sensor readout and passive RF sensors. He has authoredr co-authored more than 70 papers in international peer-reviewed journals andonference publications and he holds three patents.

ators B 199 (2014) 470–478

Dr. Angeliki Tserepi received the B.Sc. Degree in Physics from the University ofAthens, Greece, in 1985, and the M.Sc. and Ph.D. degrees from the Ohio State Uni-versity, USA, in 1989 and 1994, respectively. She held a Post-Doctoral EuropeanFellowship (“Marie Curie”) at the University Joseph Fourier of Grenoble, France,1994–1996, and at IMEL–NCSR “Demokritos”, Greece, in 1996–1997. Since 1997,she is with the Institute of Nanoscience & Nanotechnology (former IMEL), whereshe is now director of research, working on plasma processing of materials forapplications in micro- nano-technology and microsystem fabrication. Her current

international conferences. She is the author/co-author of 7 patents and more than 90publications in international peer-reviewed journals, and editor of two InternationalProceedings.