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Sensors and Actuators B 202 (2014) 60–66

Contents lists available at ScienceDirect

Sensors and Actuators B: Chemical

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

apid prototyping of multifunctional microfluidic cartridges forlectrochemical biosensing platforms

itae Kima,1, Yong Shinb,1, Simon Songa, Joohyung Leec, Jungkyu Kimc,∗

Department of Mechanical Engineering & Institute of Nano Science and Technology, Hanyang University, Seoul, Republic of KoreaInstitute of Microelectronics, A*STAR (Agency for Science, Technology and Research), 11 Science Park Road, Singapore Science Park II,ingapore 117685, SingaporeDepartment of Mechanical Engineering, Texas Tech University, Lubbock, TX 79406, USA

r t i c l e i n f o

rticle history:eceived 27 January 2014eceived in revised form 26 March 2014ccepted 5 May 2014vailable online 21 May 2014

eywords:apid prototypingorous membranene-shot valve

a b s t r a c t

A multifunctional microfluidic cartridge for electrochemical biosensing (�CEB) was developed witha cleanroom-free rapid prototyping technique. A seven-layered �CEB platform including a gold (Au)electrode substrate was fabricated by alternating patterned thin plastic and adhesive films. To provideconformal bonding between layers, pressure sensitive adhesive (PSA) tape was adopted to fabricate a leak-free, porous membranes embedded �CEB platform compatible with relatively high flow rates. In addition,the embedded porous membranes provide multiple functions in the �CEB device, including (1) control-ling fluid flow, (2) storing a dried detergent, and (3) trapping liquid waste. To demonstrate the utility ofthe �CEB, we performed an electrochemical-based immunoassay to detect various concentrations of Cre-atine Kinase (CK)-Myocardial Band (MB). During the immunoassay, the membranes controlled the flow

n-chip chemical storageicrofluidic electrochemical immunoassay

iomarker screening

path to minimize a carryover between the assay steps and released the stored dried-reagent to removenon-specifically bound detection antibodies. With a linear sweep voltammetry (LSV) electrochemicalsensing technique, a limit of detection of 0.25 ng mL−1 was achieved. The multilayered prototyping tech-nique enables rapid and low-cost fabrication of multifunctional microfluidic electrochemical devices withsingle-use sample processing components.

. Introduction

Miniaturized electrochemical sensors have been broadlyxplored for rapid and accurate detection of chemical and biologicalgents [1–5]. Integration of this technology with microfluidic sys-ems enables simple, portable and automated biosensing platforms.hese benefits make microfluidic electrochemical (EC) devices anttractive technology for point-of-care diagnostics and infectiousiseases monitoring in resource-limited settings.

To make the microfluidic EC devices more practical, a sig-ificant effort has been made to develop rapid prototypingechniques [6–10]. Soft lithography has been extensively used

ue to simplicity, low cost and versatility [7]. To fabricatehe microfluidic EC device with the soft-lithography, bondingetween a polydimethylsiloxane (PDMS) channel layer and an

∗ Corresponding author at: Department of Mechanical Engineering, Departmentf Internal Medicine, Texas Tech University, Lubbock, TX 79409, USA.el.: +1 806 834 6106.

E-mail address: jungkyu.kim@ttu.edu (J. Kim).1 These authors contributed equally to this work.

ttp://dx.doi.org/10.1016/j.snb.2014.05.009925-4005/© 2014 Elsevier B.V. All rights reserved.

© 2014 Elsevier B.V. All rights reserved.

electrode-deposited substrate was performed simply by oxygenplasma treatment [11] or physical clamping [12]. This is possibledue to the elastomeric property of PDMS which allows conformalcontact with rigid substrates such as silicon and glass. Recently,other polymeric materials like polycarbonate (PC) [13–15], poly-methyl-meta-acrylate (PMMA) [16,17], cyclic olefin copolymer(COC) [18–20] and polyimide [21] have been used as alternativesdue to undesirable properties of PDMS such as poor chemical com-patibility, gas permeability and innate hydrophobicity. Althoughthose polymers are superior to PDMS in terms of native proper-ties, fabrication, integration and packaging procedures with thosepolymers can be more difficult. For example, bonding processestypically involve the use of high temperature or solvents, resultingin compromised integrity and resolution of microstructures. As aresult, bonding of an electrode-patterned rigid polymer to anotherrigid polymer remains challenging due to leakage.

Bartholomeusz et al. [22] developed a rapid prototyping tech-nique using xurography, where a cutting plotter was used to cut

microstructures on various adhesive films (pressure sensitive andthermal activated adhesive films). Patterned adhesive films werelaminated for fabricating multilayered microfluidic devices rapidlyand inexpensively. By combining plastic microfluidic fabrication

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J. Kim et al. / Sensors and

ethods with xurography, challenges such as the fluid leakage andlectrical cross-talk issues were successfully addressed.

Membrane-integrated microfluidic chips [23–30] have beentudied for enhanced functionality and mass transport control.lthough porous membranes are used in a wide range of indus-

rial applications, their usage in microfluidic systems has beenimited to filtration [23,31] and chemical reactors [32] due to theack of simple and reliable fabrication techniques. Yager et al. [33]resented a microfluidic flow-through membrane immunoassayith dry-stored reagents in a membrane. The use of functionalizedembranes enabled the simplification of assay operation. How-

ver, on-chip integration of multiple membranes, each responsibleor a different function, remains unexplored.

In this study, we report a simple rapid prototyping of a multilay-red microfluidic cartridge for electrochemical biosensor (�CEB)ith integrated Au electrodes. By using layer-by-layer assem-

ly, the patterned plastic films and double-sided PSA tapes werelternately bonded to each other to fabricate a disposable, multi-unctional �CEB at room temperature. Three porous membranesere integrated into the �CEB during the assembly process to uti-

ize for air flow stopper, reagent storage and liquid waste absorbent.o demonstrate the utility of the �CEB, we performed a quan-itative immunoassay for CK-MB, a cardiac biomarker used foriagnosis of an acute myocardial infraction (AMI). The developmentf multilayer electrochemical biosensor from the rapid prototypeechnique, multifunctional membranes and simple electrochemicaletections will be able to facilitate advanced point-of-care deviceor various biomarker screenings.

. Experimental

.1. Fabrication of �CEB platform

Fig. 1 presents a schematic for rapid prototyping of the �CEB85 mm × 54 mm × 1.8 mm) with 3-electrode configuration whichas made of 7 patterned layers, comprised of an Au-coated PET

lm (L1), double-sided tape (L2, L4, L6), an acrylic sheet (L3), and aET film (L5, L7).

To fabricate the Au electrodes on PET film (L1), 100 nm thicku was deposited by sputtering process through a shadow mask

ig. 1. Fabrication process of a multilayered, polymeric microfluidic cartridge. (A) Explodnd double-sided tape (yellow). (B) Schematic cross-section of all layers in the stack. Liape is pierced. (C) A photograph (top view) of a 3-D microfluidic cartridge (85 mm × 54 mecond and third membrane, respectively. C, R, and W represent the counter, reference aiquid meniscus.

tors B 202 (2014) 60–66 61

containing the patterns for electrodes and alignment holes. TheAu-deposited film (L1) was then treated with plasma to removecarbon residue on the Au surfaces, and the Au electrodes fur-ther underwent an antibody immobilization process (see belowin immunoassay protocol section). The double-sided PSAs (ARseal90880, Adhesives Research), with two release liners laminated onboth sides, were patterned with a cutting plotter (FC2250-120VC,Graphtec Inc., USA) to create L2/L4/L6 layer. The acrylic sheet wascut with a CO2 laser (VLS 3.60, Universal Laser Systems) to formfluidic reservoirs in L3. Three passes of the laser beam (75% Power,20% Speed and 1000 PPI) were applied to obtain smooth cuttingedges and near-vertical sidewalls. The laser-cut acrylic sheet wascleaned in an ultrasonic water bath for 10 min, dried with N2 gas,and then annealed on a hot plate maintained at 85 ◦C for 90 min.This annealing step was included to avoid stress cracks in the acrylic[34].

Once all the layers are ready, assembly began with bonding ofthe bottom PET film (L1) and the first PSA (L2). In the first step, cut-out regions (i.e., channels and reservoirs) of the PSA were peeled offfrom the bottom release liner. An application tape, typically one-sided adhesive backed film, was applied to the peeled-off side andsqueezed to hold the pattern in place. Then the release liner wascarefully detached, leaving the engraved pattern on the applica-tion tape – pattern transfer [22]. The application tape bearing thepattern (L2) was placed on a custom-made aligner with 4 press-fitted dowel pins. Each layer has 4 alignment holes at the corners,and each hole matches the corresponding dowel pin in its positionfor accurate alignment. The PET film (L1) incorporating immobi-lized antibody was adhesively bonded to the PSA (L2) held on thealigner using a hand roller. To integrate the acrylic layer (L3) withthe PET-PSA assembly (L1/L2), the assembly was put back on thealigner with its adhesive side facing up following the removal of thetop release liner along with the application tape. The acrylic sheetwas mated with the assembly held on the aligner to form a PET-PSA-acrylic assembly (L1/L2/L3). A laser-cut absorbent pad (M3,∼790 �m thick, pure cellulose fibers, 133, Pall) was attached to an

adhesive region of the assembly (L1/L2/L3). Next, a pattern trans-ferred from the second PSA (L4) was bonded to the PET-PSA-acrylicassembly (L1/L2/L3) in the same way as described above (i.e., bond-ing of L1 and L2). The layer-by-layer assembly process was repeated

ed view of all patterned layers. Alternating layers of thin plastic film/sheet (blue)quid is drawn into an absorbent-embedded waste reservoir by vacuum when them × 1.8 mm) constructed by layer-by-layer assembly. M1, M2 and M3 are the first,nd working electrode, respectively. The sense electrodes monitor the position of a

62 J. Kim et al. / Sensors and Actua

Table 1Description of each layer regarding material, thickness, function, and patterningtool. DST represents the double-sided tape.

Material Thickness (�m) Function Pattering tool

Layer 7 PET 125 Lid Cutting plotterLayer 6 DST 130 Channel Cutting plotterLayer 5 PET 125 Support Cutting plotterLayer 4 DST 130 Via hole Cutting plotterLayer 3 Acrylic 1000 Reservoir CO laser

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Layer 2 DST 130 Channel Cutting plotterLayer 1 PET 125 Electrode Cutting plotter

ntil the top PET layer (L7) was integrated into a whole assemblyL1 through L7) as described in Fig. 1B. The whole assembly washen compressed multiple times using a desktop cold laminatoro ensure sufficient bonding strength as well as a tight seal of thentegrated Au electrodes. Table 1 shows the summary of detailednformation on each layer.

.2. Integration of porous membranes

Tween-20 detergent is routinely added to phosphate-bufferedaline (PBS) or tris-buffered saline (TBS) wash buffer to minimizehe background noise from non-specific binding in bioassays. Toimplify the on-chip assay procedure, we embedded a porous mem-rane containing dried Tween-20 in a �CEB. In this way, we couldinimize the complexity of device fabrication by eliminating the

eed for liquid buffer storage in the �CEB. At first, a porous mem-rane (∼419 �m thick, spun bonded polyester, 6613, Pall) was cut

nto 3-mm and 5-mm circles. The 3-mm disk membranes weremmersed in a tube containing 50% Tween-20 (P9416, Sigma) for

h. The wetted membranes were removed from the tube and driedn a 50 ◦C oven for 4 h to prepare detergent-modified membranes.imilarly, the 5-mm disk membranes were treated with bovineerum albumin (BSA) (A2153, Sigma). The BSA-treated membraneM1) and the Tween-20 treated membrane (M2) were then insertednto the corresponding grooves formed by three stacked layersL4/L5/L6) and tightly fixed in place when the top layer (L7) wasapped. Note that nominal thickness of the membrane was choseno be slightly higher than the depth of the groove to create tight fit.his configuration allowed liquid to block air flow when trapped inhe first membrane (M1) and to pass through the second membraneM2) for reconstitution of the detergent.

.3. Surface treatment and immunoassay preparation

To immobilize capture antibodies on the working electrode, 6 �Lf monoclonal anti-CK-MB antibody at 10 �g mL−1 (anti-humanK-MB 7501 SPRN-2, Medix Biochemica) diluted in a PBS buffer10 mM PBS, pH 7.4 (Sigma)), was dispensed onto the electrodend incubated at 4 ◦C for 2 h in a perti-dish along with a wet tissueo prevent evaporation of the antibody solution. After the incuba-ion, the electrode was washed 3 times with a TBST buffer (20 mMBS, pH 7.4 (T5912, Sigma) and 10 mg mL−1 Tween-20 (P9416,igma)). Then, TBS buffer containing 1% BSA was dropped on allhree electrodes to block the surface. After 1 h of incubation atoom temperature, the electrodes were washed 3 times with TBSTnd dried with N2 gas. Once the immobilization steps were fin-shed, the PET film (L1) was stored at 4 ◦C until used. Labeling of aetection antibody with alkaline phosphatase (ALP) was done bysing an Alkaline Phosphatase Labeling Kit-SH (LK-13-10, Dojindo,outh Korea). We followed the kit procedure to conjugate 100 �g of

etection anti-CK-MB antibody (Anti-human CK-MB 7502 SPRN-5,edix Biochemica). The conjugate solution was diluted 100 times

o be added in samples. CK-MB antigen (30-1081, Fitzgerald) waspiked and diluted to final concentrations of 1, 5, 10 and 25 ng mL−1

tors B 202 (2014) 60–66

in TBS buffer. Each concentration of the CK-MB antigen was mixedwith the ALP-labeled detection antibody at a volume ratio of 1:10and incubated at 37 ◦C for 30 min to execute the first immunore-action. The electrochemical substrate solution was prepared in amixture of 40 mM 4-aminophenyl phosphate (pAPP, A-292, Gold-bio), 1 mM MgCl2 (M2393, Sigma) and 100 mM NaCl (S3014, Sigma)in 100 mM diethyl amine (DEA, pH 9.8, 31589, Sigma).

2.4. Immunoassay for CK-MB detection

Once all the solutions were prepared, an on-chip sandwich-typeimmunoassay for CK-MB detection was performed using the �CEBplatform. Serially diluted CK-MB antigens (1, 5, 10 and 25 ng mL−1)were prepared in TBS buffer. Before starting the on-chip immunoas-say, a mixture of 45 �L of various concentrations of CK-MB antigensand 5 �L of the ALP-conjugated anti-CK-MB antibodies were incu-bated in a water bath maintained at 37 ◦C for 30 min. A 200 �Lvolume of substrate solution was pre-loaded in the substrate reser-voir (2.0 mm width × 1.2 mm height), and pierceable tape was thenattached to seal the valve’s opening (see Fig. 1B). This type of valve(normally closed) prevents a liquid plug in a conduit from movingaway from one side (valve’s position) when a negative pressure isapplied on the other side. The valve can be opened by simply per-forating the tape, resulting in the trapped air being balanced withan ambient pressure.

A 50 �L sample containing an antigen-antibody-enzyme com-plex was loaded to the sample port. The sample spontaneously filledthe sample reservoir (0.5 mm width × 1.2 mm height) by capillarywicking and stopped at the first membrane (M1). Upon applica-tion of a negative pressure via the vacuum port, the sample beganto flow into the main channel (400 �m width × 130 �m height)past the membrane to reach the absorbent pad (M3) embeddedin the waste reservoir. The flow was maintained at approximately0.21 �L s−1 for 4 min (Fig. 2A). Note that the pre-loaded substratesolution was kept in the substrate reservoir by the normally closedvalve during the sample injection. After the sample injection step,the wet membrane (M1) stops air flow from leaking into the samplereservoir.

Next, a wash step was initiated by perforating the tape witha sharp needle. The substrate solution was drawn into the mainchannel and reconstituted the dried Tween-20 in the second mem-brane (M2). Subsequently, air flow from the positive pressure portwas supplied to the junction, thereby generating air bubbles in astream of the substrate solution containing the detergent (Fig. 2B).Such a segmented flow [35] is known to assist in washing on theelectrode surface due to meniscus force and internal recirculationflow induced within liquid segments. By the time the washing stepwas finished, air bubble generation was ended by switching off thepositive pressure port (Fig. 2C). In the meantime, the substrate con-tinued flowing at a flow rate of 5.0 �L s−1 for 15 s. This is to bring afresh volume of the substrate, containing little or no Tween-20,over the electrode region. Lastly, the solution was stopped andheld immobile during the enzymatic reaction by turning off allthe pumps and then venting the remaining pressure in the �CEB(Fig. 2D). After a 5-min enzymatic reaction, linear sweep voltam-metry (LSV) signals were measured with a Parstat 2273 potentiostat(Princeton Applied Research) in the range of −0.2 to 0.3 V at a scanrate of 0.1 V s−1.

3. Results and discussion

To evaluate sealing performance of the �CEB device, we per-formed leakage tests for liquid and air in the microfluidic channels.Fig. 3A represents the magnified top view of a fluidic channel(the middle region denoted by PET) which is overlapped with Au

J. Kim et al. / Sensors and Actuators B 202 (2014) 60–66 63

Fig. 2. Images of a microfluidic cartridge during immunoassay operation. The state of pressure (positive, upper port) and vacuum (negative, lower port) pumps for each stepis denoted. (A) Sample transport to the antibody derivatized working electrode (denoted by W). Sample (red arrow) flows through the main channel for immunoreactionby turning on the negative pressure while substrate solution is held in the reservoir (orange dotted line). (B) Electrode washing and addition of substrate. Air bubbles aregenerated at the junction by applying a positive pressure, and non-specifically bound target are removed by a meniscus force (blue and orange arrows). (C) By switching offthe positive pressure, air bubbles are removed, and a fresh substrate solution was delivered over the electrodes (orange arrow). (D) By switching off all the pressure sources,the substrate solution is held on the electrodes during enzymatic reaction and electrochemical detection (orange dotted line).

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ig. 3. Leak-proof sealing of integrated electrodes using a double-sided tape. (A) A midth × 130 �m height) formed by cutting a double-sided tape. (B) A red liquid dye

t a flow rate of 9.1 �L s−1 (scale bar, 500 �m).

lectrodes on the PET film (L1). To perform a liquid leak test, a red-yed liquid was flowed through the fluidic channel and drained

nto a bottle which was connected to the waste reservoir (with nobsorbent pad in this test). The vacuum-driven flow was visuallyhecked for leakage while the flow rate was gradually increased.here was no visible leakage around the fluidic channel at flow ratesp to 9.1 �L s−1 as shown in Fig. 3B. In addition, we conducted a pre-

iminary test for gas leakage. To detect air leaks out of the �CEB, syringe pump was set to slowly increase air pressure inside the

CEB that was immersed in water. No air bubbles were observedp to 20 psi. This result demonstrates that the PSA-based bondingechnique is suitable for leak-proof sealing of metal electrodes overigid substrates.

ed view of an Au electrode-deposited PET film crossed by a fluidic channel (400 �mng in the fluidic channel which is overlapped with three electrodes (100 nm thick)

For diverse functionality of integrated membranes, we utilizedthe first membrane (M1) as a valve. As depicted in Fig. 4A, a liquidsolution continuously flows to the main channel past the M1 by avacuum pump until the solution from the reservoir is all dissipatedand captured by the membrane. At this time, a positive pressure isapplied from the right side of the main channel to allow the solu-tion to be separated from the membrane. When M1 was wetted bythe sample solution, it functions as a valve to prevent air inside thechannel from being discharged to the reservoir via the membrane.

As shown in Fig. 4B, air can only flow along the main channel butcannot flow in the direction blocked by the liquid and the mem-brane. In a similar manner, the third membrane (M3) was usedfor isolating the waste solutions from both the main channel and

64 J. Kim et al. / Sensors and Actuators B 202 (2014) 60–66

Fig. 4. Operation of the first membrane (M1) for valve function. (A) A liquid (dotted arrow) passes through the membrane and moves to the main channel by a vacuum pump.T en the liquid in the sample reservoir is all dissipated, a pressure pump applies a force tos w (blue arrow) from leaking into the sample reservoir in the air bubble generation step.

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Fig. 5. Comparison of background currents (with no target antigen) for washperformance in the sandwich electrochemical immunoassay. 100 �L samples ofelectrochemical substrate solution with (modified) vs. without (unmodified) 1%Tween-20 were applied at different flow rates of 0.71, 1.43 and 5.0 �L s−1 for thefirst wash. Subsequently, another 100 �L of unmodified substrate solution was sup-

he valve in its open state is used in the sample transport step (red arrow). (B) Wheparate the liquid from the main channel. The valve in its closed state blocks air flo

he vacuum port. This function will be very useful for the safety ofperators, particularly in screening for contagious diseases.

In addition, the second membrane (M2) was used as a driedhemical storage. Once the membrane is wetted with a buffer solu-ion, we can simply reconstitute to obtain the desired buffer. Inhe �CEB, the substrate solution was exploited for wash (modi-ed usage) as well as enzymatic reaction (original usage), greatlyimplifying an immunoassay design. For the washing step, by pass-ng the solution through a Tween-20 dried membrane, the solution

as modified to 1% Tween-20 buffer to remove the non-specificallyound protein on the electrodes. After reconstituting the driedween-20, the unmodified substrate solution was delivered overhe electrodes for enzymatic reaction. With this simplified forma-ion, wash performance during the immunoassay was evaluated byomparing non-specific background signals associated with twoypes of membranes (i.e., in the presence vs. absence of Tween-0). We used 50 �L of sample containing CK-MB antibody-enzymeonjugate and 100 �L of substrate solution (out of total 200 �L)ith or without 1% Tween-20 in the wash test. The first 100 �Las supplied at different flow rates of 0.71, 1.43 and 5.0 �L s−1, fol-

owed by the second (unmodified) 100 �L with a fixed flow ratef 5.0 �L s−1. No air bubbles were produced in this procedure inrder to only test the role of Tween-20 rehydrated from the mem-rane to washing efficiency. As shown in Fig. 5, the �CEB utilizinghe Tween-20 coated membrane yielded a lower background sig-al than one without Tween-20, demonstrating that the membraneith Tween-20 enhanced washing efficiency.

To demonstrate the assay capability of the �CEB, we performedn electrochemical immunoassay for 5 different target concen-rations (1, 5, 10, 25 ng mL−1 including a negative control). Asescribed above, a 50 �L sample containing an antigen-ALP labeledetection antibody was applied to the �CEB’s sample port, and

mmunoreaction occurred during sample flow. Then, 200 �L of sub-

trate solution was constantly pumped at a flow rate of 5.0 �L s−1

or wash and halted for a 5-min enzymatic reaction. Fig. 6A presentshe current responses at the working electrode plotted againsthe applied potential in LSV. Faradic current contribution was

plied for both second wash and enzymatic reaction. The error bar represents thestandard deviation from three independent measurements.

extracted at a fixed potential of 0.15 V to give a peak anodic cur-rent. Fig. 6B exhibits the calibration curve of the peak anodiccurrent versus the concentration of CK-MB. The current responseschanged log-linearly as the analyte concentration increased, whichindicates that the on-chip electrochemical scheme could be usedfor quantitative detection in immunoassays. From a series of theimmunoassays, we achieved a limit of detection of 0.25 ng mL−1 inthe CK-MB assay. Elevated CK-MB level has been correlated with theincidence of myocardial damage, and regular screening of CK-MBlevels has been suggested for the early detection of acute myocar-

dial infarction (AMI) [36]. Although there is no clear thresholdCK-MB value that is definitively diagnostic for AMI, ∼10 ng mL−1

has been proposed as an acceptable detection limit for early AMI

J. Kim et al. / Sensors and Actuators B 202 (2014) 60–66 65

Fig. 6. (A) Current responses obtained from on-cartridge immunoassay for CK-MB at different concentrations of 25, 10, 5, 1 and 0 ng mL−1. Linear sweep voltammetry wasr −1

procev . (B) C5 was 0

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un in the range of −0.2 to 0.3 V at a scan rate of 0.1 V s . The immunoassay wasolume of substrate solution was constantly supplied at 5.0 �L s−1 during wash step, 10 and 25 ng mL−1) including a control. The background current (with no target)

creening assays [37]. Thus, the performance of the �CEB platformor CK-MB assay is within clinical norms.

. Conclusion

The development of microfluidic systems has promoteddvances in miniaturization and automation of molecular diagnos-ic testing, however high fabrication costs are often a barrier toommercialization and/or adaptation to resource poor settings. TheCEB including a passive valve, reagent storage and electrochemi-

al sensors was fabricated by alternately stacking thin plastic filmsnd double-sided PSA tapes. Our rapid fabrication technique elim-nates the need for specialized cleanroom equipment and wouldrovide a miniaturized, affordable platform for point-of-care (POC)esting. In addition, the use of pierceable tape and porous mem-ranes in �CEB enables programmable control of fluid flow duringhe performance of an assay. This capability eliminates the need for

echanical microvalves in the design, thereby significantly reduc-ng the amount of off-chip control equipment. The process that isutomated by our �CEB platform can be useful for a broad range ofargets that employ similar immunoassay protocols. In this work,e demonstrated an on-chip electrochemical immunoassay forK-MB and achieved clinically relevant sensitivity. The rapid fab-ication technique that we present enables the design, fabrication,nd testing of new devices within a single day.

cknowledgments

This research was supported by Basic Science Research Pro-ram through the National Research Foundation of Korea (NRF)unded by the Ministry of Education, Science and Technology2012R1A6A1029029). J. Kim would like to thank Texas Techniversity for financial support of this project through new inves-

igator start-up funding. The authors would also like to thank Dr.ensen for fruitful scientific discussions.

eferences

[1] J. Kim, J. Elsnab, C. Gehrke, J. Li, B.K. Gale, Microfluidic integrated multi-walledcarbon nanotube (MWCNT) sensor for electrochemical nucleic acid concentra-tion measurement, Sens. Actuator B 185 (2013) 370–376.

[2] Y. Sameenoi, K. Koehler, J. Shapiro, K. Boonsong, Y. Sun, J. Collett Jr., et al.,Microfluidic electrochemical sensor for on-line monitoring of aerosol oxidativeactivity, J. Am. Chem. Soc. 134 (2012) 10562–10568.

[3] J. Wang, Electrochemical biosensors: towards point-of-care cancer diagnostics,Biosens. Bioelectron. 21 (2006) 1887–1892.

[

[

ssed on the microfluidic cartridge with Tween-20 modified membrane. A 200 �Lalibration plot of the anodic peak current vs. the input concentration of CK-MB (1,

.11 ± 0.014 �A. The test was repeated three times for each concentration.

[4] J.W. Choi, K.W. Oh, J.H. Thomas, W.R. Heineman, H.B. Halsall, J.H. Nevin,et al., An integrated microfluidic biochemical detection system for proteinanalysis with magnetic bead-based sampling capabilities, Lab Chip 2 (2002)27–30.

[5] A. Bange, H.B. Halsall, W.R. Heineman, Microfluidic immunosensor systems,Biosens. Bioelectron. 20 (2005) 2488–2503.

[6] T.W. de Haas, H. Fadaei, D. Sinton, Laminated thin-film Teflon chips for petro-chemical applications, Lab Chip 12 (2012) 4236–4239.

[7] D.C. Duffy, J.C. McDonald, O.J. Schueller, G.M. Whitesides, Rapid prototyp-ing of microfluidic systems in poly(dimethylsiloxane), Anal. Chem. 70 (1998)4974–4984.

[8] Y. Lu, W. Shi, L. Jiang, J. Qin, B. Lin, Rapid prototyping of paper-based microflu-idics with wax for low-cost, portable bioassay, Electrophoresis 30 (2009)1497–1500.

[9] E. Sollier, C. Murray, P. Maoddi, D. Di Carlo, Rapid prototyping polymersfor microfluidic devices and high pressure injections, Lab Chip 11 (2011)3752–3765.

10] T. Yasukawa, A. Glidle, M. Nomura, J.M. Cooper, Fabrication of robust 2-D and3-D microfluidic networks for lab-on-a-chip bioassays, J. Microelectromech.Syst. 14 (2005) 839–846.

11] A. Plecis, Y. Chen, Fabrication of microfluidic devices based onglass–PDMS–glass technology, Microelectron. Eng. 84 (2007) 1265–1269.

12] A. Yamaguchi, P. Jin, H. Tsuchiyama, T. Masuda, K. Sun, S. Matsuo, et al., Rapidfabrication of electrochemical enzyme sensor chip using polydimethylsiloxanemicrofluidic channel, Anal. Chim. Acta 468 (2002) 143–152.

13] Z. Chen, J. Wang, S. Qian, H.H. Bau, Thermally-actuated, phase change flowcontrol for microfluidic systems, Lab Chip 5 (2005) 1277–1285.

14] J. Kim, D. Byun, M.G. Mauk, H.H. Bau, A disposable, self-contained PCR chip, LabChip 9 (2009) 606–612.

15] J. Yang, Y. Liu, C.B. Rauch, R.L. Stevens, R.H. Liu, R. Lenigk, et al., High sensitivityPCR assay in plastic micro reactors, Lab Chip 2 (2002) 179–187.

16] L. Yao, B. Liu, T. Chen, S. Liu, T. Zuo, Micro flow-through PCR in a PMMA chipfabricated by KrF excimer laser, Biomed. Microdev. 7 (2005) 253–257.

17] W. Zhang, S. Lin, C. Wang, J. Hu, C. Li, Z. Zhuang, et al., PMMA/PDMS valves andpumps for disposable microfluidics, Lab Chip 9 (2009) 3088–3094.

18] Q. Pu, O. Oyesanya, B. Thompson, S. Liu, J.C. Alvarez, On-chip micropatterning ofplastic (cylic olefin copolymer, COC) microfluidic channels for the fabricationof biomolecule microarrays using photografting methods, Langmuir 23 (2007)1577–1583.

19] J. Steigert, S. Haeberle, T. Brenner, C. Muller, C.P. Steinert, P. Koltay, et al., Rapidprototyping of microfluidic chips in COC, J. Micromech. Microeng. 17 (2007)333–341.

20] R. Novak, N. Ranu, R.A. Mathies, Rapid fabrication of nickel molds forprototyping embossed plastic microfluidic devices, Lab Chip 13 (2013)1468–1471.

21] B.C. Giordano, J. Ferrance, S. Swedberg, A.F. Huhmer, J.P. Landers, Polymerasechain reaction in polymeric microchips: DNA amplification in less than 240 s,Anal. Biochem. 291 (2001) 124–132.

22] D.A. Bartholomeusz, R.W. Boutte, J.D. Andrade, Xurography rapid prototypingof microstructures using a cutting plotter, J. Microelectromech. Syst. 14 (2005)1364–1374.

23] J. Kim, B.K. Gale, Quantitative and qualitative analysis of a microfluidic DNAextraction system using a nanoporous AlOx membrane, Lab Chip 8 (2008) 1516.

24] J. Kim, M. Mauk, D. Chen, X. Qiu, J. Kim, B. Gale, et al., A PCR reactor with anintegrated alumina membrane for nucleic acid isolation, Analyst 135 (2010)2408.

25] J. Kim, R. Surapaneni, B.K. Gale, Rapid prototyping of microfluidic systems usinga PDMS/polymer tape composite, Lab Chip 9 (2009) 1290.

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6 J. Kim et al. / Sensors and

26] B.H. Chueh, D. Huh, C.R. Kyrtsos, T. Houssin, N. Futai, S. Takayama, Leakage-freebonding of porous membranes into layered microfluidic array systems, Anal.Chem. 79 (2007) 3504–3508.

27] P. Guo, E.W. Hall, R. Schirhagl, H. Mukaibo, C.R. Martin, R.N. Zare, Microfluidiccapture and release of bacteria in a conical nanopore array, Lab Chip 12 (2012)558–561.

28] F. Saharil, F. Forsberg, Y. Liu, P. Bettotti, N. Kumar, F. Niklaus, et al., Dry adhesivebonding of nanoporous inorganic membranes to microfluidic devices using theOSTE(+) dual-cure polymer, J. Micromech. Microeng. 23 (2013) 025021.

29] F. Tan, P.H.M. Leung, Z.-b. Liu, Y. Zhang, L. Xiao, W. Ye, et al., A PDMS microflu-idic impedance immunosensor for E. coli O157:H7 and Staphylococcus aureusdetection via antibody-immobilized nanoporous membrane, Sens. Actuator B159 (2011) 328–335.

30] J. Yu, Z. Liu, Q. Liu, K.T. Yuen, A.F.T. Mak, M. Yang, et al., A polyethylene glycol(PEG) microfluidic chip with nanostructures for bacteria rapid patterning anddetection, Sens. Actuator A 154 (2009) 288–294.

31] H.B. Wei, B.H. Chueh, H.L. Wu, E.W. Hall, C.W. Li, R. Schirhagl, et al., Particlesorting using a porous membrane in a microfluidic device, Lab Chip 11 (2011)238–245.

32] D.S. Peterson, T. Rohr, F. Svec, J.M.J. Frechet, Enzymatic microreactor-on-a-chip: protein mapping using trypsin immobilized on porous polymer monolithsmolded in channels of microfluidic devices, Anal. Chem. 74 (2002) 4081–4088.

33] D.Y. Stevens, C.R. Petri, J.L. Osborn, P. Spicar-Mihalic, K.G. McKenzie, P. Yager,Enabling a microfluidic immunoassay for the developing world by integrationof on-card dry reagent storage, Lab Chip 8 (2008) 2038–2045.

34] H. Klank, J.P. Kutter, O. Geschke, CO2-laser micromachining and back-endprocessing for rapid production of PMMA-based microfluidic systems, Lab Chip2 (2002) 242–246.

35] V. Linder, S.K. Sia, G.M. Whitesides, Reagent-loaded cartridges for valveless andautomated fluid delivery in microfluidic devices, Anal. Chem. 77 (2005) 64–71.

36] D. Chan, L.L. Ng, Biomarkers in acute myocardial infarction, BMC Med. 8 (2010)34.

37] A. Qureshi, Y. Gurbuz, J.H. Niazi, Biosensors for cardiac biomarkers detection:a review, Sens. Actuator B 171 (2012) 62–76.

iographies

itae Kim received a Ph.D. in mechanical & aerospace engineering from the Uni-ersity of California, Irvine (2006). His B.S. and M.S. degrees, both in mechanical

tors B 202 (2014) 60–66

engineering, are from Hanyang University in South Korea (1999) and the Universityof Southern California (2001), respectively. He is currently a research professor atthe Institute of Nano Science & Technology (INST), Hanyang University. His interestsand research include droplet-based microfluidics, point-of-care (POC) immunoassayand microfluidic nucleic acid testing (NAT).

Yong Shin received the M.S. degree in Cancer Biology from Seoul National Univer-sity, South Korea in 2005 and Ph.D. degree in Molecular Neuro-Biology from MaxPlanck Institute of Experimental Medicine and Georg-August-University Goettin-gen, Germany in 2008. He currently joined Institute of Microelectronics, A*STAR asa Scientist, Singapore. His research is now focused on the development of moleculardiagnostic platform based on optical biophotonics for detection of disease relatedbiomarkers.

Simon Song received his B.S. (1995) in Mechanical Engineering Department,Hanyang University, Seoul, Korea and M.S. (1997) and Ph.D. (2002) in Mechan-ical Engineering Department, Stanford University, USA. He joined faculties atHanyang University since 2004 and has interest in the development of microflu-idic chips for chemical sensing or sensor synthesis on a microchip as well asflow visualization technology using medical equipment like magnetic resonanceimaging.

Joohyung Lee received both B.S. (2009) and M.S. (2011) degrees in Biomedical Engi-neering from Yonsei University, South Korea. He is currently a Ph.D. student in thebiomedical micro/nano device (BMND) lab at the Texas Tech University, USA. Hisresearch activities are now focused on developing microengineered biomimeticorgan-on-a-chip and chemical/biochemical analysis platforms for environmentalmonitoring and biomarker screening.

Jungkyu (Jay) Kim is Assistant Professor, Department of Mechanical Engi-neering at Texas Tech University. His research expertise includes lab-on-achip devices that require a variety of microfluidic components for com-plex chemical/biomedical assays. His work has involved the development ofmicrofluidic control system, microfluidic sample processing, nucleic acid sam-

ple preparation and on-chip amplification, protein microarray, CNT biosensorand immunomagnetic detection. He has authored and co-authored more than50 peer-reviewed journal and conference publications, 1 book chapters, and 6patents issued or pending in the area of microfluidics, biosensor and cell/tissueengineering.