The authors acknowledge the Robert W. Deutsch Foundation, the Defense Threat Reduction Agency (DTRA), and the National Science Foundation Emerging Frontiers in Research and Innovation (EFRI) for financial support. The authors also thank the Maryland Nanocenter and its Fablab for cleanroom facility support.

1. Microfabrication of the Microfluidic Chip
<ol>
	<li>Prepare Electrochemical Chip
	<ol>
		<li>Pattern Gold Electrodes
		<ol>
			<li>Rinse a blank 4&#8217;&#8217; silicon wafer (prime grade quality) with acetone, methanol, and isopropanol (&#8220;AMI&#8221; clean). Rinse the isopropanol from the wafer with deionized water (DI) followed by drying with N2 gun.</li>
			<li>Grow a 1 &#181;m thick SiO2 passivation layer with plasma-enhanced chemical vapor deposition (PECVD) tool. Deposit a 200 &#197; thick layer of chromium followed by a 2,000 &#197; thick layer of gold with a DC sputtering tool.</li>
			<li>Spin &#8220;PR1&#8221; positive photoresist (spinning parameters: 3,000 rpm, 1,000 rpm/sec, 30 sec) to create uniform film ~1.6 &#181;m thick. Pre-bake the wafer for 1 min at 100 &#176;C on a hot plate.</li>
			<li>Expose the wafer with 190 mJ/cm2 at 405 nm wavelength through a photomask. Develop the wafer for 30 sec in 352 developer and rinse the wafer with DI and dry with N2 gun.</li>
			<li>Etch the gold with gold etchant for 1.5 min. Etch the chromium with chromium etchant for ~30 sec. Rinse the wafer with DI and dry with N2 gun.</li>
			<li>Strip the photoresist with &#8220;AMI&#8221; clean procedure. 
			NOTE: STRONG OXIDIZER AND POTENTIALLY EXPLOSIVE.</li>
			<li>Clean the wafer with 4:1 H2SO4:H2O2 mixture (&#8220;Piranha&#8221; clean) for 1 min. Rinse the wafer with DI and dry with N2 gun.</li>
		</ol>
		</li>
		<li>Pattern Platinum Electrodes
		<ol>
			<li>Spin &#8220;PR2&#8221; positive photoresist (spinning parameters: 3,000 rpm, 1,000 rpm/sec, 30 sec) to create uniform film ~1.4 &#181;m thick. Pre-bake the wafer for 1 min at 100 &#176;C on a hot plate.</li>
			<li>Expose the wafer with 30 mJ/cm2 at 405 nm wavelength through a photomask. Bake the wafer for 45 sec at 125 &#176;C on a hot plate. Flood expose the wafer with 1,000 mJ/cm2 at 405 nm wavelength without a photomask. Develop the wafer for 2 min in 6:1 AZ400K developer and rinse the wafer with DI and dry with N2 gun.</li>
			<li>Deposit a 400 &#197; thick layer of titanium followed by a 1,600 &#197; thick layer of platinum using an E-beam evaporation system.</li>
			<li>Lift off photoresist in an ultrasonicated acetone bath for 5 min followed by rinse with acetone and isopropanol. Rinse the wafer with DI and dry with N2 gun.</li>
		</ol>
		</li>
	</ol>
	</li>
	<li>Prepare Assay Channels
	<ol>
		<li>Prepare Mold for Assay Channels
		<ol>
			<li>Rinse a blank 4&#8217;&#8217; silicon wafer (test grade quality) with &#8220;AMI&#8221; clean procedure. Rinse the isopropanol from the wafer with DI followed by drying with N2 gun.</li>
			<li>Spin &#8220;PR3&#8221; negative photoresist (2-step spinning parameters: step one - 600 rpm, 120 rpm/sec, 10 sec. Step two &#8211; 1,150 rpm, 383 rpm/sec, 27 sec) to create uniform film ~100 &#181;m thick. Pre-bake the wafer on a hot plate (2-step baking parameters: step one &#8211; ramp up from room temperature to 65 &#176;C at 300 &#176;C/hr and hold temperature for 10 min. Step 2 &#8211; ramp up to 95 &#176;C at 300 &#176;C/hr and hold temperature for 30 min).</li>
			<li>Expose the wafer with 2,500 mJ/cm2 at 405 nm wavelength through a photomask. Post-bake the wafer by ramping up to 95 &#176;C at 300 &#176;C/hr and hold temperature for 10 min on a hot plate. Develop the wafer for 10 min in Propylene Glycol Monomethyl Ether Acetate (PGMEA) developer. Rinse the wafer with isopropanol and dry with N2 gun.</li>
		</ol>
		</li>
		<li>Fabricate Assay Channels
		<ol>
			<li>Prepare 10:1 polydimethylsiloxane (PDMS) mixture (20 g of elastomer, 2 g of curing agent) and completely degas in a vacuum chamber for 20 min.</li>
			<li>Define an enclosure around the mold wafer with aluminum foil and slowly pour the uncured PDMS over the mold.</li>
			<li>Bake the mold with the PDMS in a box furnace (2-step baking parameters: step one &#8211; 5 min ramp up from room temperature to 80 &#176;C. Step 2 &#8211; hold at 80 &#176;C for 17 min).</li>
			<li>Peel away the PDMS from the mold and place on aluminum foil. Cut the assay chips with a surgeon knife, define holes with a biopsy punching tool.</li>
		</ol>
		</li>
	</ol>
	</li>
</ol>
2. Assemble the Device
<ol>
	<li>Manually align the PDMS assay channels with the working electrodes on the electrochemical chip. PDMS sticks well to the electrochemical chip, sealing the channels and preventing leakage.</li>
	<li>Connect a flexible tubing (OD 0.087&#8217;&#8217; and ID 0.015&#8217;&#8217;) to a plastic elbow or straight connector using a short flexible tube as an adapter (OD 0.1875&#8217;&#8217; and ID 0.0625&#8217;&#8217;). Attach connectors to each of the hole-punched inlets in the device.</li>
	<li>Connect a syringe on the other end of the tubing and place the syringe in a syringe pump. Alternately change the set of a syringe, a tube, and a connector according to the required procedure step in section 3 and its corresponding solution.</li>
</ol>
3. Analyze DNA Hybridization
NOTE: Introduce each solution slowly into the channel at a flow rate of 200 &#181;l/hr until the whole channel is filled.
<ol>
	<li>Functionalize each vertical microchannel with a different ssDNA probe sequence (perform in parallel to the 3 separate vertical microchannels):
	<ol>
		<li>Fill each microchannel with a different ssDNA probe incubation solution (10 mM phosphate buffer, 100 mM NaCl, 10 &#181;M tris(2-carboxyethyl)phosphine (TCEP), and 1 &#181;M probe ssDNA), and incubate for 1 hr followed by rinsing the microchannel with PBS.</li>
		<li>Fill the microchannel with a solution containing 1 mM of 6-mercapto-1-hexanol (MCH) in a buffer of 10 mM PBS, 100 mM NaCl and 1.395 mM TCEP, and incubate for 1 hr followed by rinsing the microchannel with PBS.</li>
	</ol>
	</li>
	<li>Lift off the PDMS, rotate 90&#176; to a horizontal orientation, and place down to expose separate rows of reaction chambers with each row containing a unique counter and reference electrode (Figure S1).</li>
	<li>Fill the microchannels with control measurement solution of 4x concentrated saline-sodium citrate (SSC) buffer, and incubate for 20 min.</li>
	<li>Background measurement: Fill the microchannels with 5 mM ferricyanide / 5 mM ferrocyanide in PBS solution and record the impedance values of the electrochemical system for different frequencies (electrochemical impedance spectroscopy; EIS) using a potentiostat connected to the working, counter, and reference electrodes (EIS parameters: frequency range from 1 MHz to 0.1 Hz, 10 frequency data points per decade, 25 mV amplitude, 5 mV polarization versus the reference electrode). Repeat this measurement for each working electrode in the microchannels.</li>
	<li>Measure DNA hybridization (perform for each non-complementary and complementary ssDNA target).
	<ol>
		<li>Fill the microchannels with target ssDNA measurement solution of 4x concentrated SSC buffer containing 1 &#181;M of the target ssDNA, and incubate for 20 min. Measure the DNA according to step 3.4.</li>
	</ol>
	</li>
</ol>

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The authors have nothing to disclose.

Our procedures demonstrate the manufacturing of a microfluidic-based electrochemical biochip and its utilization for DNA hybridization events analysis. Through a high yield controlled microfabrication process we develop a device comprised of microscale channels integrated with an array of electrochemical transducers. We have devised controlled processing parameters for the photolithography procedure of the electrochemical chip and the microfluidic channel mold through an iterative approach. These steps provide guidelines...

Microfluidic lab-on-a-chip (LOC) devices provide numerous advantages in clinical diagnostics, environmental monitoring and biomedical research. These devices utilize microfluidic channels to control fluid flow to regions of the chip where a variety of procedures can take place including reagent mixing, affinity based binding, signal transduction, and cell culturing1-4. Microfluidics provides many advantages over conventional clinical diagnostic tools such as microwell plate readers or electrophoretic gel shift assays. Microfluidic devices require 2 to 3 orders of magnitude (nanoliters as opposed to microliters) fewer reagents to perform similar assays. Also, these devices can increase the speed by which some biological events occur due to the smaller confinement of the species within the channels5,6. Thirdly, sensors can be integrated within microfluidic devices using lithography and etching techniques, which can provide label-free detection. Lastly, these devices are inexpensive to produce and require little work on the part of the technician to operate7-10.
Label-free detection typically is performed using an optical or electrical transducer. Optical devices can present better sensing performance due to lower interference with analytes in the sample. Nevertheless, their performance is compromised in cases where the sample&#8217;s background has the same resonating wavelength as the sensor11. There are many advantages to using electrical signals to perform biological and chemical detection in microfluidic systems. The fabrication is inherently less complicated since these sensors typically only require patterned electrodes to operate. In addition, electrical signals can be directly interfaced with most measurement equipment while other signal modalities may require a transducer to convert the signal12-15. Electrical sensors commonly measure changes in impedance16,17, capacitance18, or redox activity19. However, new challenges are presented as these systems are miniaturized. The most important challenges to overcome include: sample preparation and mixing of fluids (due to the low sample volume and Reynolds number), physical and chemical effects (including capillary forces, surface roughness, chemical interactions between construction materials and analytes), low signal-to-noise ratio (produced by the reduced surface area and volume)20-23, and potential interference from electro-active analytes in complex biological samples (e.g., blood and saliva). Further investigation of these effects will result in guidelines for an accurate fabrication and operation of these devices in a reproducible manner that would improve upon their overall performance.
DNA hybridization detection is used extensively to diagnose genetic disorders24,25 and various forms of cancer26. Every year, multiple strains of influenza are identified in patients using results from DNA hybridization techniques27. The influenza virus alone accounts for 36,000 deaths each year in the United States28. Such examples could benefit from a bench-top microfluidic device that can perform the same assay techniques as a plate reader or gel shift assay with low sample volume and at a fraction of the cost without sacrificing sensitivity or specificity. Due to the many advantages of label-free electrochemical sensing, it has been used extensively for detection of DNA hybridization events29,30. A setup where macro-scale electrodes (in the millimeter range) are dipped in beakers with the solution of interest can be used to provide very sensitive data regarding the binding kinetics of single stranded DNA sequences to their matching complementary sequences. Recently, there have been a few advances in incorporating electrochemical sensing in microfluidics for DNA hybridization. Studies have been performed regarding hybridization kinetics31 and integration of sensors for detection15 in microfluidic channels. However, there still exists a need for a rapid high throughput microfluidic device that can analyze DNA hybridization events in parallel without complicated sample preparation steps.
The device presented in this work provides a platform that allows for multiple interactions to be screened in parallel and without complicated sample preparation steps. Our protocol presents how microfluidic-based electrochemical biochips are microfabricated with micro-electromechanical systems (MEMS) technology32,33. We describe the fabrication process of both the microfluidic chip, made of polydimethylsiloxane (PDMS), and the electrochemical chip, comprised of an array of electrodes. The chemical functionalization of the biochip with ssDNA probes is also addressed. Finally, the ability of the biosensor to specifically detect and analyze ssDNA targets is demonstrated. Overall, the microfluidic-based electrochemical biochip is a rapid and high-throughput analysis technique. It can be used to investigate interactions between biological molecules and conducting transducers, and can be utilized in a variety of lab-on-a-chip applications.

<table border="0" cellpadding="0" cellspacing="0" width="765">
	<tbody>
		<tr height="42">
			<td height="42" style="height:42px;width:216px;">Name of Material/ Equipment</td>
			<td style="width:155px;">Company</td>
			<td style="width:227px;">Catalog Number</td>
			<td style="width:167px;">Comments/Description</td>
		</tr>
		<tr height="21">
			<td height="21" style="height:21px;width:216px;">Material</td>
			<td style="width:155px;"></td>
			<td style="width:227px;"></td>
			<td></td>
		</tr>
		<tr height="21">
			<td height="21" style="height:21px;width:216px;">4&#39;&#39; Silicon wafer</td>
			<td style="width:155px;">Ultrasil</td>
			<td style="width:227px;">4-5664</td>
			<td>Single side polished; P-type; Boron doped; Orientation 1-0-0; Thickness 500 micron; Resistance 10-20 ohm*cm</td>
		</tr>
		<tr height="42">
			<td height="42" style="height:42px;width:216px;">Shipley 1813 photoresist (&#34;PR1&#34;)</td>
			<td style="width:155px;">Microchem</td>
			<td style="width:227px;"></td>
			<td style="width:167px;">positive photoresist</td>
		</tr>
		<tr height="21">
			<td height="21" style="height:21px;width:216px;">AZ5214 photoresist (&#34;PR2&#34;)</td>
			<td style="width:155px;">Hoechst Celanse</td>
			<td style="width:227px;"></td>
			<td style="width:167px;">positive photoresist</td>
		</tr>
		<tr height="21">
			<td height="21" style="height:21px;width:216px;">SU-8 50 photoresist (&#34;PR3&#34;)</td>
			<td style="width:155px;">Microchem</td>
			<td style="width:227px;"></td>
			<td style="width:167px;">negative photoresist</td>
		</tr>
		<tr height="21">
			<td height="21" style="height:21px;width:216px;">Gold etchant</td>
			<td style="width:155px;">Transene</td>
			<td style="width:227px;">TFA</td>
			<td style="width:167px;">No dilution</td>
		</tr>
		<tr height="21">
			<td height="21" style="height:21px;width:216px;">Chromium etchant</td>
			<td style="width:155px;">Transene</td>
			<td style="width:227px;">1020AC</td>
			<td style="width:167px;">No dilution</td>
		</tr>
		<tr height="21">
			<td height="21" style="height:21px;width:216px;">PDMS elastomer</td>
			<td style="width:155px;">Dow Corning</td>
			<td style="width:227px;">3097366 1004</td>
			<td></td>
		</tr>
		<tr height="21">
			<td height="21" style="height:21px;width:216px;">PDMS curing agent</td>
			<td style="width:155px;">Dow Corning</td>
			<td style="width:227px;">3097358 1004</td>
			<td></td>
		</tr>
		<tr height="21">
			<td height="21" style="height:21px;width:216px;">Biopsy punching tool</td>
			<td style="width:155px;">Healthlink</td>
			<td style="width:227px;">BP20</td>
			<td></td>
		</tr>
		<tr height="21">
			<td height="21" style="height:21px;width:216px;">Tygon flexible tubing</td>
			<td style="width:155px;">Cole Parmer&#160;</td>
			<td style="width:227px;">R 3603</td>
			<td>.015&#34;</td>
		</tr>
		<tr height="21">
			<td height="21" style="height:21px;width:216px;">Tygon flexible adapter</td>
			<td style="width:155px;">Cole Parmer&#160;</td>
			<td style="width:227px;">06417-41</td>
			<td>.0625&#34;</td>
		</tr>
		<tr height="21">
			<td height="21" style="height:21px;width:216px;">1 mL syringe</td>
			<td style="width:155px;">Beckton-Dickenson</td>
			<td align="right" style="width:227px;">301025</td>
			<td></td>
		</tr>
		<tr height="42">
			<td height="42" style="height:42px;width:216px;">Monobasic potassium phosphate</td>
			<td style="width:155px;">Fluka</td>
			<td align="right" style="width:227px;">1551139</td>
			<td></td>
		</tr>
		<tr height="42">
			<td height="42" style="height:42px;width:216px;">Potassium phosphate dibasic anhydrous</td>
			<td style="width:155px;">Sigma</td>
			<td style="width:227px;">RES20765-A7</td>
			<td></td>
		</tr>
		<tr height="21">
			<td height="21" style="height:21px;width:216px;">Sodium chloride</td>
			<td style="width:155px;">Sigma-Aldrich</td>
			<td style="width:227px;">S7653</td>
			<td></td>
		</tr>
		<tr height="21">
			<td height="21" style="height:21px;width:216px;">tris(2-carboxyethyl)phosphine</td>
			<td style="width:155px;">Aldrich</td>
			<td style="width:227px;">C4706</td>
			<td></td>
		</tr>
		<tr height="21">
			<td height="21" style="height:21px;width:216px;">6-mercapto-1-hexanol</td>
			<td style="width:155px;">Aldrich</td>
			<td align="right" style="width:227px;">451088</td>
			<td></td>
		</tr>
		<tr height="42">
			<td height="42" style="height:42px;width:216px;">20x concentrated saline-sodium citrate buffer</td>
			<td style="width:155px;">Sigma</td>
			<td align="right" style="width:227px;">93017</td>
			<td></td>
		</tr>
		<tr height="42">
			<td height="42" style="height:42px;width:216px;">Potassium hexacyanoferrate(III)</td>
			<td style="width:155px;">Aldrich</td>
			<td align="right" style="width:227px;">455946</td>
			<td></td>
		</tr>
		<tr height="21">
			<td height="21" style="height:21px;width:216px;">Sodium ferrocyanide</td>
			<td style="width:155px;">Aldrich</td>
			<td style="width:227px;">CDS001589</td>
			<td></td>
		</tr>
		<tr height="42">
			<td height="42" style="height:42px;width:216px;">Target ssDNA #1</td>
			<td style="width:155px;">Integrated DNA Technologies (IDT)</td>
			<td style="width:227px;">100 nmole DNA oligo</td>
			<td>5&#8217;-/5ThioMC6-D/AAAGCTCCGATAGCGCTCCG 
			TGGACGTCCC-3&#8217;</td>
		</tr>
		<tr height="42">
			<td height="42" style="height:42px;width:216px;">Complementary ssDNA #1</td>
			<td style="width:155px;">Integrated DNA Technologies (IDT)</td>
			<td style="width:227px;">100 nmole DNA oligo</td>
			<td>5&#8217;-GGGACGTCCACGGAGCGCTA 
			TCGGAGCTTT-3&#8217;</td>
		</tr>
		<tr height="42">
			<td height="42" style="height:42px;width:216px;">Target ssDNA #2</td>
			<td style="width:155px;">Integrated DNA Technologies (IDT)</td>
			<td style="width:227px;">100 nmole DNA oligo</td>
			<td>5&#8217;-/5ThioMC6-D/ACGCGTCAGGTCATTGACGA 
			ATCGATGAGT-3&#8217;</td>
		</tr>
		<tr height="42">
			<td height="42" style="height:42px;width:216px;">Complementary ssDNA #2</td>
			<td style="width:155px;">Integrated DNA Technologies (IDT)</td>
			<td style="width:227px;">100 nmole DNA oligo</td>
			<td>5&#8217;-ACTCATCGATTCGTCAATGA 
			CCTGACCCGT-3&#8217;</td>
		</tr>
		<tr height="42">
			<td height="42" style="height:42px;width:216px;">Target ssDNA #3</td>
			<td style="width:155px;">Integrated DNA Technologies (IDT)</td>
			<td style="width:227px;">100 nmole DNA oligo</td>
			<td>5&#8217;-/5ThioMC6-D/ACCTAGATCCAGTAGTTAGA 
			CCCATGATGA-3&#8217;</td>
		</tr>
		<tr height="42">
			<td height="42" style="height:42px;width:216px;">Complementary ssDNA #3</td>
			<td style="width:155px;">Integrated DNA Technologies (IDT)</td>
			<td style="width:227px;">100 nmole DNA oligo</td>
			<td>5&#8217;-TCATCATGGGTCTAACTACT 
			GGATCTAGGT-3&#8217;</td>
		</tr>
		<tr height="21">
			<td height="21" style="height:21px;width:216px;"></td>
			<td style="width:155px;"></td>
			<td style="width:227px;"></td>
			<td></td>
		</tr>
		<tr height="21">
			<td height="21" style="height:21px;width:216px;">Instrument</td>
			<td style="width:155px;"></td>
			<td style="width:227px;"></td>
			<td></td>
		</tr>
		<tr height="42">
			<td height="42" style="height:42px;width:216px;">Plasma Enhanced Chemical Vapor Deposition (PECVD)</td>
			<td style="width:155px;">Oxford Instruments</td>
			<td style="width:227px;">PlasmaLab System 100</td>
			<td>Chamber pressure 1000 mTorr, Tempreture 200 celsius degrees, RF power 20 W, Gasses: a) N2O flow rate 710 sccm; b) a composition of 5% SiH4 and 95% N2 gasses flow rate 170 sccm, SiO2 growth rate 690 Angstrom/minute.</td>
		</tr>
		<tr height="21">
			<td height="21" style="height:21px;width:216px;">DC sputtering unit</td>
			<td style="width:155px;">AJA International</td>
			<td style="width:227px;">ATC 1800-V</td>
			<td>For Chrome: Chamber pressure 10 mTorr, Argon flow rate 20 sccm, Supplied DC power 200 W, Sputter rate 10 nm/min. For Gold: Chamber pressure 10 mTorr, Argon flow rate 20 sccm, Supplied DC power 200 W, Sputter rate 36 nm/min.</td>
		</tr>
		<tr height="21">
			<td height="21" style="height:21px;width:216px;">E-beam evaporation system</td>
			<td style="width:155px;">Denton</td>
			<td style="width:227px;">Custom-built</td>
			<td>For Titanium: Chamber pressure &#60;2E-6 Torr, Supplied power to the e-gun 7.5k W, Filament current 150-200 mA, Evaporation rate 6-10 Angstrom/second. For Platinum: Chamber pressure &#60;2E-6 Torr, Supplied power to the e-gun 7.5k W, Filament current 150-200 mA, Evaporation rate 2-3 Angstrom/second.&#160;</td>
		</tr>
		<tr height="21">
			<td height="21" style="height:21px;width:216px;">Potentiostat</td>
			<td style="width:155px;">CH Instruments</td>
			<td style="width:227px;">660D</td>
			<td></td>
		</tr>
		<tr height="21">
			<td height="21" style="height:21px;width:216px;">Syringe pump</td>
			<td style="width:155px;">KD Scientific</td>
			<td style="width:227px;">KDS230</td>
			<td></td>
		</tr>
	</tbody>
</table>

a microfluidic-based electrochemical biochip for label-free dna hybridization analysis

Miniaturization of analytical benchtop procedures into the micro-scale provides significant advantages in regards to reaction time, cost, and integration of pre-processing steps. Utilizing these devices towards the analysis of DNA hybridization events is important because it offers a technology for real time assessment of biomarkers at the point-of-care for various diseases. However, when the device footprint decreases the dominance of various physical phenomena increases. These phenomena influence the fabrication precision and operation reliability of the device. Therefore, there is a great need to accurately fabricate and operate these devices in a reproducible manner in order to improve the overall performance. Here, we describe the protocols and the methods used for the fabrication and the operation of a microfluidic-based electrochemical biochip for accurate analysis of DNA hybridization events. The biochip is composed of two parts: a microfluidic chip with three parallel micro-channels made of polydimethylsiloxane (PDMS), and a 3 x 3 arrayed electrochemical micro-chip. The DNA hybridization events are detected using electrochemical impedance spectroscopy (EIS) analysis. The EIS analysis enables monitoring variations of the properties of the electrochemical system that are dominant at these length scales. With the ability to monitor changes of both charge transfer and diffusional resistance with the biosensor, we demonstrate the selectivity to complementary ssDNA targets, a calculated detection limit of 3.8 nM, and a 13% cross-reactivity with other non-complementary ssDNA following 20 min&#160;of incubation. This methodology can improve the performance of miniaturized devices by elucidating on the behavior of diffusion at the micro-scale regime and by enabling the study of DNA hybridization events.

A controllable and accurate manufacturing process for the experimental device is essential in research. It allows researchers to obtain reproducible and high throughput experiments. Here we have demonstrated a high yield, high reproducibility microfabrication process of a microfluidic-based electrochemical biochip (Figure 1). With a low failure rate, few devices have shown bonding issues that lead to solution leakage. In order to validate the electrochemical activity of the biochip, cyclic voltammetry me...

Watch this Scientific Journal Video about A Microfluidic-based Electrochemical Biochip for Label-free DNA Hybridization Analysis at JoVE.com

A Microfluidic-based Electrochemical Biochip for Label-free DNA Hybridization Analysis

We present a microfluidic-based electrochemical biochip for DNA hybridization detection. Following ssDNA probe functionalization, the specificity, sensitivity, and detection limit are studied with complementary and non-complementary ssDNA targets. Results illustrate the influence of the DNA hybridization events on the electrochemical system, with a detection limit of 3.8 nM.

Miniaturization of analytical benchtop procedures into the micro-scale provides significant advantages in regards to reaction time, cost, and integration ...

a-microfluidic-based-electrochemical-biochip-for-label-free-dna

Research

JoVE Journal

Bioengineering

17.0K Views.  University of Maryland. We present a microfluidic-based electrochemical biochip for DNA hybridization detection. Following ssDNA probe functionalization, the specificity, sensitivity, and detection limit are studied with complementary and non-complementary ssDNA targets. Results illustrate the influence of the DNA hybridization events on the electrochemical system, with a detection limit of 3.8 nM.

Video: A Microfluidic-based Electrochemical Biochip for Label-free DNA Hybridization Analysis

A Microfluidic-based Hydrodynamic Trap for Single Particles

The ability to confine and manipulate single particles in free solution is a key enabling technology for fundamental and applied science. Methods for particle trapping based on optical, magnetic, electrokinetic, and acoustic techniques have led to major advancements in physics and biology ranging from the molecular to cellular level. In this article, we introduce a new microfluidic-based technique for particle trapping and manipulation based solely on hydrodynamic fluid flow. Using this method, we demonstrate trapping of micro- and nano-scale particles in aqueous solutions for long time scales. The hydrodynamic trap consists of an integrated microfluidic device with a cross-slot channel geometry where two opposing laminar streams converge, thereby generating a planar extensional flow with a fluid stagnation point (zero-velocity point). In this device, particles are confined at the trap center by active control of the flow field to maintain particle position at the fluid stagnation point. In this manner, particles are effectively trapped in free solution using a feedback control algorithm implemented with a custom-built LabVIEW code. The control algorithm consists of image acquisition for a particle in the microfluidic device, followed by particle tracking, determination of particle centroid position, and active adjustment of fluid flow by regulating the pressure applied to an on-chip pneumatic valve using a pressure regulator. In this way, the on-chip dynamic metering valve functions to regulate the relative flow rates in the outlet channels, thereby enabling fine-scale control of stagnation point position and particle trapping. The microfluidic-based hydrodynamic trap exhibits several advantages as a method for particle trapping. Hydrodynamic trapping is possible for any arbitrary particle without specific requirements on the physical or chemical properties of the trapped object. In addition, hydrodynamic trapping enables confinement of a "single" target object in concentrated or crowded particle suspensions, which is difficult using alternative force field-based trapping methods. The hydrodynamic trap is user-friendly, straightforward to implement and may be added to existing microfluidic devices to facilitate trapping and long-time analysis of particles. Overall, the hydrodynamic trap is a new platform for confinement, micromanipulation, and observation of particles without surface immobilization and eliminates the need for potentially perturbative optical, magnetic, and electric fields in the free-solution trapping of small particles.

In this article, we present a microfluidic-based method for particle confinement based on hydrodynamic flow. We demonstrate stable particle trapping at a fluid stagnation point using a feedback control mechanism, thereby enabling confinement and micromanipulation of arbitrary particles in an integrated microdevice.

The ability to confine and manipulate single particles in free solution is a key enabling technology for fundamental and applied science.  Methods for ...

Fabrication of Electrochemical-DNA Biosensors for the Reagentless Detection of Nucleic Acids, Proteins and Small Molecules

As medicine is currently practiced, doctors send specimens to a central laboratory for testing and thus must wait hours or days to 
receive the results. Many patients would be better served by rapid, bedside tests. To this end our laboratory and others have developed a versatile, reagentless 
biosensor platform that supports the quantitative, reagentless, electrochemical detection of nucleic acids (DNA, RNA), proteins (including antibodies) and small 
molecules analytes directly in unprocessed clinical and environmental samples. In this video, we demonstrate the preparation and use of several biosensors in this 
"E-DNA" class. In particular, we fabricate and demonstrate sensors for the detection of a target DNA sequence in a polymerase chain reaction mixture, an HIV-specific antibody and the drug cocaine. The preparation procedure requires only three hours of hands-on effort followed by an overnight incubation, and their use requires only minutes.

"E-DNA" sensors, reagentless, electrochemical biosensors that perform well even when challenged directly in blood and other complex matrices, have been adapted to the detection of a wide range of nucleic acid, protein and small molecule analytes. Here we present a general procedure for the fabrication and use of such sensors.

As medicine is currently practiced, doctors send specimens to a central laboratory for testing and thus must wait hours or days to 
receive the results. ...

Non-contact, Label-free Monitoring of Cells and Extracellular Matrix using Raman Spectroscopy

Non-destructive, non-contact and label-free technologies to monitor cell and tissue cultures are needed in the field of biomedical research.1-5 However, currently available routine methods require processing steps and alter sample integrity. Raman spectroscopy is a fast method that enables the measurement of biological samples without the need for further processing steps. This laser-based technology detects the inelastic scattering of monochromatic light.6 As every chemical vibration is assigned to a specific Raman band (wavenumber in cm-1), each biological sample features a typical spectral pattern due to their inherent biochemical composition.7-9 Within Raman spectra, the peak intensities correlate with the amount of the present molecular bonds.1 Similarities and differences of the spectral data sets can be detected by employing a multivariate analysis (e.g. principal component analysis (PCA)).10
Here, we perform Raman spectroscopy of living cells and native tissues. Cells are either seeded on glass bottom dishes or kept in suspension under normal cell culture conditions (37 &deg;C, 5% CO2) before measurement. Native tissues are dissected and stored in phosphate buffered saline (PBS) at 4 &deg;C prior measurements. Depending on our experimental set up, we then either focused on the cell nucleus or extracellular matrix (ECM) proteins such as elastin and collagen. For all studies, a minimum of 30 cells or 30 random points of interest within the ECM are measured. Data processing steps included background subtraction and normalization.

Raman spectroscopy is a suitable technique for the non-contact, label-free analysis of living cells, tissue-engineered constructs and native tissues. Source-specific spectral fingerprints can be generated and analyzed using multivariate analysis.

Non-destructive, non-contact and label-free technologies to monitor cell and tissue cultures are needed in the field of biomedical research.1-5 However, ...

Hollow Microneedle-based Sensor for Multiplexed Transdermal Electrochemical Sensing

The development of a minimally invasive multiplexed monitoring system for rapid analysis of biologically-relevant molecules could offer individuals suffering from chronic medical conditions facile assessment of their immediate physiological state. Furthermore, it could serve as a research tool for analysis of complex, multifactorial medical conditions. In order for such a multianalyte sensor to be realized, it must be minimally invasive, sampling of interstitial fluid must occur without pain or harm to the user, and analysis must be rapid as well as selective.
Initially developed for pain-free drug delivery, microneedles have been used to deliver vaccines and pharmacologic agents (e.g., insulin) through the skin.1-2 Since these devices access the interstitial space, microneedles that are integrated with microelectrodes can be used as transdermal electrochemical sensors. Selective detection of glucose, glutamate, lactate, hydrogen peroxide, and ascorbic acid has been demonstrated using integrated microneedle-electrode devices with carbon fibers, modified carbon pastes, and platinum-coated polymer microneedles serving as transducing elements.3-7,8
This microneedle sensor technology has enabled a novel and sophisticated analytical approach for in situ and simultaneous detection of multiple analytes. Multiplexing offers the possibility of monitoring complex microenvironments, which are otherwise difficult to characterize in a rapid and minimally invasive manner. For example, this technology could be utilized for simultaneous monitoring of extracellular levels of, glucose, lactate and pH,9 which are important metabolic indicators of disease states7,10-14 (e.g., cancer proliferation) and exercise-induced acidosis.15

This article details the construction of a multiplexed microneedle-based sensor. The device is being developed for in situ sampling and electrochemical analysis of multiple analytes in a rapid and selective manner. We envision clinical medicine and biomedical research uses for these microneedle-based sensors.

The development of a minimally invasive multiplexed monitoring system for rapid analysis of biologically-relevant molecules could offer individuals ...

Bacterial Detection &amp; Identification Using Electrochemical Sensors

Electrochemical sensors are widely used for rapid and accurate measurement of blood glucose and can be adapted for detection of a wide variety of analytes. Electrochemical sensors operate by transducing a biological recognition event into a useful electrical signal. Signal transduction occurs by coupling the activity of a redox enzyme to an amperometric electrode. Sensor specificity is either an inherent characteristic of the enzyme, glucose oxidase in the case of a glucose sensor, or a product of linkage between the enzyme and an antibody or probe.
Here, we describe an electrochemical sensor assay method to directly detect and identify bacteria. In every case, the probes described here are DNA oligonucleotides. This method is based on sandwich hybridization of capture and detector probes with target ribosomal RNA (rRNA). The capture probe is anchored to the sensor surface, while the detector probe is linked to horseradish peroxidase (HRP). When a substrate such as 3,3',5,5'-tetramethylbenzidine (TMB) is added to an electrode with capture-target-detector complexes bound to its surface, the substrate is oxidized by HRP and reduced by the working electrode. This redox cycle results in shuttling of electrons by the substrate from the electrode to HRP, producing current flow in the electrode.

We describe an electrochemical sensor assay method for rapid bacterial detection and identification. The assay involves a sensor array functionalized with DNA oligonucleotide capture probes for ribosomal RNA (rRNA) species-specific sequences. Sandwich hybridization of target rRNA with the capture probe and a horseradish peroxidase-linked DNA oligonucleotide detector probe produces a measurable amperometric current.

Electrochemical  sensors are widely used for rapid and accurate measurement of blood glucose and  can be adapted for detection of a wide variety of ...

Multi-analyte Biochip (MAB) Based on All-solid-state Ion-selective Electrodes (ASSISE) for Physiological Research

Lab-on-a-chip (LOC) applications in environmental, biomedical, agricultural, biological, and spaceflight research require an ion-selective electrode (ISE) that can withstand prolonged storage in complex biological media 1-4. An all-solid-state ion-selective-electrode (ASSISE) is especially attractive for the aforementioned applications. The electrode should have the following favorable characteristics: easy construction, low maintenance, and (potential for) miniaturization, allowing for batch processing. A microfabricated ASSISE intended for quantifying H+, Ca2+, and CO32- ions was constructed. It consists of a noble-metal electrode layer (i.e. Pt), a transduction layer, and an ion-selective membrane (ISM) layer. The transduction layer functions to transduce the concentration-dependent chemical potential of the ion-selective membrane into a measurable electrical signal.
The lifetime of an ASSISE is found to depend on maintaining the potential at the conductive layer/membrane interface 5-7. To extend the ASSISE working lifetime and thereby maintain stable potentials at the interfacial layers, we utilized the conductive polymer (CP) poly(3,4-ethylenedioxythiophene) (PEDOT) 7-9 in place of silver/silver chloride (Ag/AgCl) as the transducer layer. We constructed the ASSISE in a lab-on-a-chip format, which we called the multi-analyte biochip (MAB) (Figure 1).
Calibrations in test solutions demonstrated that the MAB can monitor pH (operational range pH 4-9), CO32- (measured range 0.01 mM - 1 mM), and Ca2+ (log-linear range 0.01 mM to 1 mM). The MAB for pH provides a near-Nernstian slope response after almost one month storage in algal medium. The carbonate biochips show a potentiometric profile similar to that of a conventional ion-selective electrode. Physiological measurements were employed to monitor biological activity of the model system, the microalga Chlorella vulgaris.
The MAB conveys an advantage in size, versatility, and multiplexed analyte sensing capability, making it applicable to many confined monitoring situations, on Earth or in space.
Biochip Design and Experimental Methods
The biochip is 10 x 11 mm in dimension and has 9 ASSISEs designated as working electrodes (WEs) and 5 Ag/AgCl reference electrodes (REs). Each working electrode (WE) is 240 &mu;m in diameter and is equally spaced at 1.4 mm from the REs, which are 480 &mu;m in diameter. These electrodes are connected to electrical contact pads with a dimension of 0.5 mm x 0.5 mm. The schematic is shown in Figure 2.
Cyclic voltammetry (CV) and galvanostatic deposition methods are used to electropolymerize the PEDOT films using a Bioanalytical Systems Inc. (BASI) C3 cell stand (Figure 3). The counter-ion for the PEDOT film is tailored to suit the analyte ion of interest. A PEDOT with poly(styrenesulfonate) counter ion (PEDOT/PSS) is utilized for H+ and CO32-, while one with sulphate (added to the solution as CaSO4) is utilized for Ca2+. The electrochemical properties of the PEDOT-coated WE is analyzed using CVs in redox-active solution (i.e. 2 mM potassium ferricyanide (K3Fe(CN)6)). Based on the CV profile, Randles-Sevcik analysis was used to determine the effective surface area 10. Spin-coating at 1,500 rpm is used to cast ~2 &mu;m thick ion-selective membranes (ISMs) on the MAB working electrodes (WEs).
The MAB is contained in a microfluidic flow-cell chamber filled with a 150 &mu;l volume of algal medium; the contact pads are electrically connected to the BASI system (Figure 4). The photosynthetic activity of Chlorella vulgaris is monitored in ambient light and dark conditions.

Multi-analyte Biochip MAB Based on All-solid-state Ion-selective Electrodes ASSISE for Physiological Research

All-solid-state ion-selective electrodes (ASSISEs) constructed from a conductive polymer (CP) transducer provide several months of functional lifetime in liquid media. Here, we describe the fabrication and calibration process of ASSISEs in a lab-on-a-chip format. The ASSISE is demonstrated to have maintained a near-Nernstian slope profile after prolonged storage in complex biological media.

Lab-on-a-chip (LOC) applications in environmental, biomedical, agricultural, biological, and spaceflight research require an ion-selective electrode (ISE) ...

Microfluidic-based Electrotaxis for On-demand Quantitative Analysis of Caenorhabditis elegans' Locomotion

The nematode Caenorhabditis elegans is a versatile model organism for biomedical research because of its conservation of disease-related genes and pathways as well as its ease of cultivation. Several C. elegans disease models have been reported, including neurodegenerative disorders such as Parkinson's disease (PD), which involves the degeneration of dopaminergic (DA) neurons 1. Both transgenes and neurotoxic chemicals have been used to induce DA neurodegeneration and consequent movement defects in worms, allowing for investigations into the basis of neurodegeneration and screens for neuroprotective genes and compounds 2,3.
Screens in lower eukaryotes like C. elegans provide an efficient and economical means to identify compounds and genes affecting neuronal signaling. Conventional screens are typically performed manually and scored by visual inspection; consequently, they are time-consuming and prone to human errors. Additionally, most focus on cellular level analysis while ignoring locomotion, which is an especially important parameter for movement disorders.
We have developed a novel microfluidic screening system (Figure 1) that controls and quantifies C. elegans' locomotion using electric field stimuli inside microchannels. We have shown that a Direct Current (DC) field can robustly induce on-demand locomotion towards the cathode ("electrotaxis") 4. Reversing the field's polarity causes the worm to quickly reverse its direction as well. We have also shown that defects in dopaminergic and other sensory neurons alter the swimming response 5. Therefore, abnormalities in neuronal signaling can be determined using locomotion as a read-out. The movement response can be accurately quantified using a range of parameters such as swimming speed, body bending frequency and reversal time.
Our work has revealed that the electrotactic response varies with age. Specifically, young adults respond to a lower range of electric fields and move faster compared to larvae 4. These findings led us to design a new microfluidic device to passively sort worms by age and phenotype 6.
We have also tested the response of worms to pulsed DC and Alternating Current (AC) electric fields. Pulsed DC fields of various duty cycles effectively generated electrotaxis in both C. elegans and its cousin C. briggsae 7. In another experiment, symmetrical AC fields with frequencies ranging from 1 Hz to 3 KHz immobilized worms inside the channel 8.
Implementation of the electric field in a microfluidic environment enables rapid and automated execution of the electrotaxis assay. This approach promises to facilitate high-throughput genetic and chemical screens for factors affecting neuronal function and viability.

Microfluidic-based Electrotaxis for On-demand Quantitative Analysis of Caenorhabditis elegans&#039; Locomotion

A semi-automated micro-electro-fluidic method to induce on-demand locomotion in Caenorhabditis elegans is described. This method is based on the neurophysiologic phenomenon of worms responding to mild electric fields (&#x201C;electrotaxis&#x201D;) inside microfluidic channels. Microfluidic electrotaxis serves as a rapid, sensitive, low-cost, and scalable technique to screen for factors affecting neuronal health.

The nematode Caenorhabditis elegans is a versatile model organism for biomedical research because of its conservation of disease-related genes and ...

Use of Label-free Optical Biosensors to Detect Modulation of Potassium Channels by G-protein Coupled Receptors

Ion channels control the electrical properties of neurons and other excitable cell types by selectively allowing ions to flow through the plasma membrane1. To regulate neuronal excitability, the biophysical properties of ion channels are modified by signaling proteins and molecules, which often bind to the channels themselves to form a heteromeric channel complex2,3. Traditional assays examining the interaction between channels and regulatory proteins require exogenous labels that can potentially alter the protein&#39;s behavior and decrease the physiological relevance of the target, while providing little information on the time course of interactions in living cells. Optical biosensors, such as the X-BODY Biosciences BIND Scanner system, use a novel label-free technology, resonance wavelength grating (RWG) optical biosensors, to detect changes in resonant reflected light near the biosensor. This assay allows the detection of the relative change in mass within the bottom portion of living cells adherent to the biosensor surface resulting from ligand induced changes in cell adhesion and spreading, toxicity, proliferation, and changes in protein-protein interactions near the plasma membrane. RWG optical biosensors have been used to detect changes in mass near the plasma membrane of cells following activation of G protein-coupled receptors (GPCRs), receptor tyrosine kinases, and other cell surface receptors. Ligand-induced changes in ion channel-protein interactions can also be studied using this assay. In this paper, we will describe the experimental procedure used to detect the modulation of Slack-B sodium-activated potassium (KNa) channels by GPCRs.

Optical biosensor techniques can detect changes in mass near the plasma membrane in living cells and allow one to follow cellular responses in both individual cells and populations of cells. This protocol will describe detection of the modulation of potassium channels by G-protein coupled receptors in intact cells using this approach.

Ion channels control the electrical properties of neurons and other excitable cell types by selectively allowing ions to flow through the plasma ...

A Label-free Technique for the Spatio-temporal Imaging of Single Cell Secretions

Inter-cellular communication is an integral part of a complex system that helps in maintaining basic cellular activities. As a result, the malfunctioning of such signaling can lead to many disorders. To understand cell-to-cell signaling, it is essential to study the spatial and temporal nature of the secreted molecules from the cell without disturbing the local environment. Various assays have been developed to study protein secretion, however, these methods are typically based on fluorescent probes which disrupt the relevant signaling pathways. To overcome this limitation, a label-free technique is required.
In this paper, we describe the fabrication and application of a label-free localized surface plasmon resonance imaging (LSPRi) technology capable of detecting protein secretions from a single cell. The plasmonic nanostructures are lithographically patterned onto a standard glass coverslip and can be excited using visible light on commercially available light microscopes. Only a small fraction of the coverslip is covered by the nanostructures and hence this technique is well suited for combining common techniques such as fluorescence and bright-field imaging.
A multidisciplinary approach is used in this protocol which incorporates sensor nanofabrication and subsequent biofunctionalization, binding kinetics characterization of ligand and analyte, the integration of the chip and live cells, and the analysis of the measured signal. As a whole, this technology enables a general label-free approach towards mapping cellular secretions and correlating them with the responses of nearby cells.

Inter-cellular communication is critical for controlling various physiological activities within and outside the cell. This paper describes a protocol for measuring the spatio-temporal nature of single cell secretions. To achieve this, a multidisciplinary approach is used which integrates label-free nanoplasmonic sensing with live cell imaging.

Inter-cellular communication is an integral part of a complex system that helps in maintaining basic cellular activities. As a result, the malfunctioning ...

Label-free Single Molecule Detection Using Microtoroid Optical Resonators

Detecting small concentrations of molecules down to the single molecule limit has impact on areas such as early detection of disease, and fundamental studies on the behavior of molecules. Single molecule detection techniques commonly utilize labels such as fluorescent tags or quantum dots, however, labels are not always available, increase cost and complexity, and can perturb the events being studied. Optical resonators have emerged as a promising means to detect single molecules without the use of labels. Currently the smallest particle detected by a non-plasmonically-enhanced bare optical resonator system in solution is a 25 nm polystyrene sphere1. We have developed a technique known as Frequency Locking Optical Whispering Evanescent Resonator (FLOWER) that can surpass this limit and achieve label-free single molecule detection in aqueous solution2. As signal strength scales with particle volume, our work represents a &#62; 100x improvement in the signal to noise ratio (SNR) over the current state of the art. Here the procedures behind FLOWER are presented in an effort to increase its usage in the field.

We have developed a label-free biosensing system based on optical resonator technology known as Frequency Locking Optical Whispering Evanescent Resonator (FLOWER) that is capable of detecting single molecules in solution. Here the procedures behind this work are described and presented.

Detecting small concentrations of molecules down to the single molecule limit has impact on areas such as early detection of disease, and fundamental ...