Biomolecular Sensing Processing and Analysis - Rashid Bashir and Steve Wereley

146	ABRAHAM P. LEE, JOHN COLLINS, AND ASUNCION V. LEMOFF

The work done in moving a charge from point a to point b is simply the force multiplied by the distance w given by:

Wab = Fw.

(7.10)

Substituting the Lorentz force into Eq. 7.10 results in:

Wab = qν Bw,

(7.11)

where q is the charge of the particle, ν is the velocity of particle and B the magnetic field. The voltage difference is defined as the work per unit charge which is given by:

Vab = ν Bw.			(7.12)
The flow velocity is related to the volumetric flow rate Q by:
ν =	Q		(7.13)
		,
	wh

where wh is the cross-sectional area of the channel. Substituting Eq. 7.13 into 7.12 gives:

Vab =	Q B	(7.14)
	h .

Thus, measuring the voltage difference across the channels gives us the flow rate.

7.3.2. MHD Based Viscosity meter

A viscosity meter is simply a pump and a flow meter in series with one another. The first set of electrodes is used as the pump and the second set of electrodes as a flow meter. Measuring the voltage difference from the flow meter allows us to determine the flow rate. For the MHD pump, the flow rate for a rectangular cross-section channel is given by Eq. 7.6. Since I, B, w, h, and L are known, from the pump, and Q is measured by the flow-meter, the viscosity µ, can be deduced. Viscosity measurements are important in µTAS since biological particles in a solution affect the viscosity of the solution.

7.3.3.Impedance Sensors with MicroChannel Parallel Electrodes

7.3.3.1.Electrical Double Layer In a microchannel with a pair of electrodes interfacing the liquid the electrostatic charges on the electrode surface will attract the counter ions in the liquid. The rearrangement of the charges on the electrode surface balances the charges in the liquid. This gives rise to electrical double layer [10] (EDL). Because of the electrostatic interaction, the counter ion concentration near the electrode surface is higher than that in the bulk liquid far away from the solid surface. Immediately next to the solid surface, there is a layer of ions that are strongly attracted to the solid surface and are immobile. This compact layer, normally less than 1 nm thick is called Helmholtz layer, labeled as inner (IHP) and outer (OHP) Helmholtz plane in Fig. 7.11. From the compact layer to the uniform bulk liquid, the counter ion concentration gradually reduces to that of bulk liquid. Ions in this region are affected less by the electrostatic interaction and are mobile. This layer

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potential

electrode

OH Diffusion

IH

FIGURE 7.11. Electrical Double Layer.

is called the diffuse layer of the EDL as proposed by Gouy and Chapman. The thickness of the diffuse layer depends on the bulk ionic concentration and electrical properties of the liquid, ranging from a few nanometers for high ionic concentration solutions up to 1 mm for distilled water and pure organic liquids. Stern put together compact layer and diffusion layer while Grahame proposed the possibility of some ionic or uncharged species to penetrate in to the zone closest to the electrodes. The potential rises rapid when it moves from Helmholtz plane to diffusion layer and goes to saturation in the middle of the channel. Thus, the flow of liquids along a pair of electrodes reorients the ionic distribution in the channel.

7.3.3.2. Fabrication of Channel Electrodes and Microfluidic Channel The impedance sensors were generally fabricated with thin surface metallic electrodes, fabricated [32] on a glass substrate (Fig. 7.12) and the microfluidic channel being

FIGURE 7.12. (a) Layout design of fabricated electrodes and the wiring for measuring the current flow.

(b) Cross-section of fabricated flow sensor along the electrodes at AA’ in (a).

148	ABRAHAM P. LEE, JOHN COLLINS, AND ASUNCION V. LEMOFF

Outlet

Flow

Inlet

AC

FIGURE 7.13. Measurements using Channel Electrodes.

made on polydimethylsiloxane (PDMS) using a SU8 mold [8]. Electrodes have also been electroplated for 3-d electrodes which results in uniform electric fields for higher sensitivity and accuracy [20, 21].

7.3.3.3. Flow Sensing When a liquid is forced through a microchannel under an applied hydrostatic pressure, the counter ions in the diffuse layer of the EDL are carried toward the downstream end, resulting in an electrical current in the pressure-driven flow direction. This current is called the streaming current. However, this current is measured with two electrodes on the line of direction of flow through the channel. On the other hand, in order to use the channel electrodes for measuring flow impedance or admittance measurements across the channel electrodes is considered. An application of an ac signal across the electrodes results in an increase in electrical admittance across the electrodes. This increase of admittance increases with flow rate of the liquid flow through the channel. Thus the channel electrodes act as a flow sensor in the channel with the measurement of flow induced admittance [7] (see Fig. 7.13).

In hydrodynamic conditions, forced convection dominates the transport of ions to the electrodes within the flow channels. When the width of the microfluidic channel is very small compared to the length of the channel, the lateral diffusion of the ions is significant under laminar flow. Under an ac electrical signal applied across the channel, the equivalent circuit [28] of the microsystem is shown in Fig. 7.14a. The electrical double layer [10] formed across the channel is formed from two capacitances namely diffuse layer capacitance (Cs) and the outer Helmholtz plane capacitance (Ce). The former is due to ion excess or depletion in the channel, and the latter is due to the free electrons at the electrodes and is independent of the electrolyte concentration. The smaller of these capacitances dominates the admittance since these two capacitances are in series. The frequency of the applied ac voltage,

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FIGURE 7.14. (a) The equivalent circuit for the channel and electrodes flow sensor cell. The solution in the channel offers a parallel resistive (Rs) and capacitive (Cs) impedance while the electrodes by themselves offer serial capacitive (Ce) impedance with the solution. (b) Experimental Setup for measuring current increase due to flow of electrolytes (Standard Resistance R = 1k , AC is the ac signal source, NI DAQ is PCI 6024E).

flow rate and conductivity of the fluid are the factors affecting the admittance of the fluidic system and our flow sensing principle is based on the optimization of these parameters.

∂2[ A]					∂[ A]		= 0		(7.15)
DA			− νx
	∂ y2				∂ x
		∂ A		= 0					(7.16)
		∂t
iL = 0.925n F [ A]bulk D2A/3 Q1/3w · 3									(7.17)
								xe2/ h2d

For an electrochemical oxidation of a species A to A+ in a microchannel, the convectivediffusive equation for mass transport under steady state condition is given by equation (1), where [A] is the concentration of the species, DA is the diffusion coefficient and vx is the velocity in the direction of flow. The first term is the lateral diffusion in the microchannel and the second term is the transport along the length of the channel. Under steady state flow condition the boundary condition is given by equation (2). The solution of this equation predicts the mass transport limited current, iL [18] as a function of flow rate, Q as given by equation (3), where n is the number of electrons transferred, F, the Faraday constant, xe is the electrode length, h, the cell half-height, d, the width of the cell and w, the electrode width. It is to be noted that the current due to flow of electrolyte is directly proportional to the cube root of volume flow rate of the fluid. AC voltage signal is considered rather than dc voltage since the application of an ac voltage in the flow sensor does not promote any electrode reaction.

In order to measure the flow induced admittance an ac voltage is applied across the channel electrodes in series with a standard resistor and the current flowing across the pair of electrodes is calculated. The voltage across the resistor is fed to a data acquisition card through an amplifier as shown in Fig. 14b. The rms values of voltage across the standard resistor are measured using a programmable interface to the computer.

Microfluidic flow of an electrolyte (Eg. NaOH) is maintained at a constant flow rate in the channel using a syringe pump. An ac signal of low voltage (Eg. 0.05 V) is applied in the circuit by a signal generator. The current exponentially grows when the flow is switched on and then stays constant. After switching off the flow, the current again decays exponentially until it reaches a constant value. The difference between two constant values of current gives the current increase due to flow and this current is a key figure in the measurements.

The flow sensor is optimized for the electrical parameters (f = 500 Hz, V = 0.4 V and concentration = 0.2 M) and the flow sensor is capable of measuring very low values

150	ABRAHAM P. LEE, JOHN COLLINS, AND ASUNCION V. LEMOFF

	6.0x10-4						0.6
(mA)						0.4
	-4	a
	4.0x10			0.2
Current
		0.05		0.1
	2.0x10-4
	0.0
	0	100	200	300	400	500	600

Time (sec)

Current (mA)

						0.8
3.0x10-4	b					0.6
2.0x10-4						0.4
						0.2
1.0x10-4						0.0
0.0	0.1	0.2	0.3	0.4	0.5	0.6
		Flow Rate (uL/min)

(mm/sec) Velocity

FIGURE 7.15. Flow sensor calibration (a) Time versus current across the sensor electrodes when flow of fluid is turned on for 1 minute and turned off for 1 minutes at the flow rates 0.05 µL/min to 0.6 µL/min (b) Flow induced currents calculated from (a) and bead velocities at different flow rates.

of flow rate starting at 0.05 µL/min (< 1 nL/sec). The current measured is proportional to the flow rate as shown in Fig. 7.15. In another experiment, fluorescent beads of diameter 2.5 µm are mixed with NaOH and sent through the channel. The motion of the beads at the flow rates .05, 0.1, 0.2, 0.4, and 0.6 µL/min are recorded using optical video microscopy at 30, 60, 120, 250, 250 frames/sec respectively, and the beads at a particular stream are analyzed and averaged to predict the velocity response at very low flow rates. The sensor results are compared with the velocity of beads (the symbol ‘ ’ in Fig. 7.15b) and shows similar response. Thus the calibration of the flow sensor is accomplished using the velocity measurements with beads.

7.3.3.4. Measurement of Solution Properties Microfabricated impedance sensors [3] have demonstrated the ability to sense variations in solution temperatures, ionic concentrations, and even antigen-antibody binding (immunosensors) in microchannels. Traditional admittance spectra of solution represent different ionic dispersions at broad frequency range. It is very difficult to quantitatively analyse the spectra of different solutions. On the other hand, the flow of a solution along the impedance sensing channel electrodes in the flow induced admittance measurement configuration gives more information on the solution. The flow induced admittance depends on the frequency of the ac signal applied across the electrodes for the admittance measurement. The flow induced admittance frequency spectra of different fluids flowing across the channel is characteristic of the molecules or constituents of the fluid.

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Flow Induced Admittance

0.003	PBS		0.0020				D-MEM
0.002			0.0015
0.001			0.0010
0.000
500	1000	1500	2000	100	200	300	400	500
0.003			0.003
		KOH					NaOH

0.002

0.002

					0.001
0.001					0.000
200		400	600	800	1000		0		500	1000	1500	2000

Frequency (Hz)

FIGURE 7.16. Flow induced admittance spectra of some analytical solutions.

Fig. 7.16 shows the flow induced admittance spectra for cell culture reagents PBS and D-MEM and ionic solutions KOH and NaOH. These spectra for different solutions not only differ in magnitude but also show clear shift in frequency and width of the peak. The peak magnitude, frequency and band width of the spectra depend upon the ionic properties like ionic strength, valency of the ions, concentration of the ions, etc, present in the solutions.

7.3.3.5. Measurement of Particles in Solution Measurement electrical properties of tiny particles suspended in solution or any liquid have been measured in their bulk form using microelectrodes are electrodes arrays built in a well. With the advent of microfluidics, a single particle is pumped across electrical sensors based on capacitance or impedance measurement and are sensed. Thus a solution containing non-identical particles can be sensed continuously with the pair of electrodes in the channel. These sensors, not only can count the particles based on their electrical response, but also can detect the nature of the particles based on the relationship between the electrical parameters and the nature of the particle.

Particle sensing has been carried out traditionally using fluorescence or radioactive tagging. Measurements based on such methods are very sensitive but that require optical staining or radioactive labelling and other manipulations. Electrical measurements are better than such techniques in the sense that they do not require sophisticated sample preparation techniques.

Depending on the electrical behaviour of the liquids and the particles, capacitance or impedance measurements are employed. Generally, capacitance measurements are sensitive

152	ABRAHAM P. LEE, JOHN COLLINS, AND ASUNCION V. LEMOFF

1.0E+9
1.0E+8
		Chromafin Cells
1.0E+7		Red Blood Cells
1.0E+6		Chromafin Cells
1.0E+5
1.0E+4
1.0E+3
1.0E+2	Red Blood Cells

1.0E+1

1.0E+0

0.0E+0 5.0E+5 1.0E+6 1.5E+6 2.0E+6 2.5E+6 3.0E+6 3.5E+6 Frequency (Hz)

Phase

5.0E+0

0.0E	+	0	5.0E+6 1.0E+6 1.5E+6 2.0E+6 2.5E+6 3.0E+6 3.5E+6
-5.0E+0
-1.5E+1
-2.5E+1				Chromafin Cells
-3.5E+1
-4.5E+1
-5.5E+1				Red Blood Cells
-6.5E+1
-7.5E+1
-8.5E+1					Chromafin Cells
					RBS
-9.5E+1			Frequency (Hz)

FIGURE 7.17. Single cell impedance spectroscopy of bovine chromaffin and red blood cells [21].

at low frequencies and impedance measurements are sensitive at high frequencies. If the liquid (pH buffer, saline solution) where the particles are suspended is more conducting than the particles dielectric of the particles are predominant and so capacitance measurements are more sensitive. On the other hand if the liquid is a non-conductor (eg. DI water, oil, solvent) impedance measurements can detect the particles suspended in the liquid. In a typical microfluidic sensor, a pair of electrodes is built across the channel where the particles are flowed in the channel. The length of the electrodes is comparable to the size of the particles to be measured. In order to make sure that the particles are flowing one by one, the width of the channel is less than twice the size of the particles.

Single cell characterization [21] of bovine chromaffin cells and red blood cells has been conducted using electrical impedance spectroscopy over a frequency range of 40 Hz to 3 MHz. In order to trap the cell in between opposing electrodes, vacuum or pressure and dielectrophoresis techniques have been utilized. The impedance measurements are done for the cell media in order to provide baseline for the impedance data recorded for the cells along with the media. At lower frequencies of the impedance spectra, large difference in cell impedance were observed whereas a characteristic impedance value develops at higher frequencies due to the elimination of the membrane capacitive component. The phase data is very sensitive for different cells types than the magnitude spectrum of the cells as shown in Fig. 7.17.

7.3.3.6. Particle Cytometry using Capacitance and Impedance Measurements In cytometry, particles are sensed one by one continuously so that monitoring of every particle is possible. Micro Coulter particle counter principle is utilized in most of the cytometries. Capacitance cytometry [27] is based on ac capacitance measurement by probing the polarization response of the particles in an external electric field. Capacitance measurements have been used to assay cell-cycle progression [1], differentiate normal and malignant white blood cells [24], DNA content within the nucleus of the cell [27], cell growth etc. Capacitance cytometry of cells is carried using the channel electrodes by measuring the capacitance of the cells when they flow through the channel.

The integrated microfluidic device for the cell cytometry consists of a pair of electrodes where the cells are sensed, and the PDMS microfluidic channel with inlet and outlet holes for fluid.

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FIGURE 7.18. Capacitance Cytometry of mouse cells correlated to the DNA contents of the cells under different metabolism cycles [27].

The electrodes are made of gold and are 50 µm wide. The distance, d, separating the electrodes is 30 µm. The width of the PDMS microfluidic channel is also d, the length, L, is 5 mm, and the height, h, is either 30 µm or 40 µm. Fluid delivery is accomplished with a syringe pump at nonpulsatile rates ranging from 1 to 300 µl/hr. By electrically shielding the device and controlling the temperature precisely (to within ±0.05◦C), noise levels of ≈5 aF when the microfluidic channel is dry and 0.1-2 fF when wet, are achieved.

Figure 7.18 shows the device response over a course of 1,000 ms to fixed mouse myeloma SP2/0 cells suspended in 75% ethanol/25% PBS solution at 10◦C. Distinct peaks present in the data correspond to a single cell flowing past the electrodes. The channel height of the device was 30 µm. The peak values of the capacitance are correlated to the DNA content of the different cells and also to the metabolism phase cycle of the cells.

7.4. PARALLEL MICROCHANNEL ELECTRODES FOR SAMPLE PREPARATION

7.4.1.A Microfluidic Electrostatic DNA Extractor

7.4.1.1.DNA Extractor Principle DNA is captured and concentrated electrostatically using a microfluidic device that utilizes the inherent negative charge on a DNA molecule for its capture. Due to the advances in molecular biology, techniques for reading the genome (DNA sequencing) and for identifying the existence of a known sequence (DNA detection) have been developed. However, extracting the DNA from a raw sample, such as a blood or urine sample, can be labor-intensive and time consuming. For blood samples, DNA extraction involves several steps by a trained technician. These steps involve measuring the sample volume, cell sorting using a centrifuge, lysing the cell (breaking down the cell membrane), and filtering out the DNA for detection. Each of the steps mentioned involves a carefully controlled set of procedures in order to be done correctly. DNA extractor chip

154	ABRAHAM P. LEE, JOHN COLLINS, AND ASUNCION V. LEMOFF

has been developed by Cepheid using silicon pillars [14]. The silicon pillars are oxidized which allow for DNA to bind to the surface as biological samples are pumped through. Once the DNA has been concentrated, a wash solution followed by a buffer solution is flowed through the microstructure to release the DNA.

We present another method to extract DNA from a sample using the inherent net negative charge on the DNA. The DNA Extractor described uses the H-filterTM design developed by Micronics Technologies combined with electrostatic forces. The principle of the H-filterTM relies on the absence of turbulent mixing in a microfluidic channel. Thus, two flow streams can flow next to one another without mixing. Movement of particles from one flow stream to another occurs due to diffusion coefficient, particle size, viscosity of the solution and temperature.

7.4.1.2. DNA Extractor Design and Experiment DNA molecules have a net charge of 2 negative charges per base pair. Thus, in the presence of a DC electric field, DNA molecules are attracted to the positive electrode. Using the H-filterTM design, DNA extraction can be achieved by patterning electrodes along the bar of the H-filterTM and applying a DC voltage across the channel. Instead of relying on diffusion to extract DNA from one flow stream to another, electrostatic forces are used to transport DNA from one stream to another.

The DNA extractor can be used to remove the DNA from a lysed spore or cell for example. The lysing solution breaks down the spore coating or cell membrane. This allows the DNA to be transported from the lysing solution stream to a buffer solution. The extraction not only allows DNA concentration but also allows the DNA to be separated from other debris that could inhibit DNA detection techniques, such as PCR (polymerase chain reaction). This is illustrated in Fig. 7.19.

The DNA extractor has two sample inputs. The first sample input is the sample volume containing the DNA. The test sample volume consists of 250 µl of DNA (48 kbp Lambda DNA from Sigma) in concentrations of 500 µg/ml of deionized water, stained with 25 µl of 1 mM YOYO-1 dye (from Molecular Probes) diluted 100:1 in distilled water and 25 µl

Buffer solution	DNA sample for further
input	processing or detection

+V

-V

DNA molecules

Lysing solution

Buffer solution

DNA sample in	To waste
lysing solution input

+V

-V

FIGURE 7.19. (a) DNA extractor chip based on the H-filterTM. DNA in a lysing solution migrates toward the positive electrode. (b) The light band along the upper half of the device is fluorescence, which indicates the presence of DNA only in the upper output channel [15].

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of 0.05% Tween 20, a surfactant to prevent sticking of the DNA to the glass/electrode surface.

The second sample input consists of 25 mM NaCl solution. An external syringe pump was used to provide flow. The infusion flow rate was set at 0.6 ml/hr. An epi-fluorescent microscope with a CCD camera is positioned above the device to view DNA migration. DNA transport is observed when the fluorescence moves from one flow stream to another.

7.4.1.3. DNA Extractor Demonstration Measurements were done to determine what DC voltage would result in electrolysis. For our channel width and exposed electrode area, 1 V was used. Any voltage above 1 V resulted in bubbles which impeded flow for both fluid streams. When the two input solutions are flowed through the device with no electrode voltage, there is sufficient diffusion over the length of the electrode, that at the output end of the electrode, DNA is present throughout the full width of the channel. Consequently, fluorescence is observed in both of the output channels, indicating the presence of DNA flowing to both outputs. When the experiment is repeated with the electrode voltage turned on, fluorescence is only observed in the output channel corresponding to the +V electrode. This is shown quite clearly in Fig. 7.19b, which clearly demonstrates the utility of this device for DNA concentration.

As can be seen in Fig. 7.19b, when the device is operating with voltage on, the DNA is concentrated sufficiently close to the +V electrode so as to be completely obscured by the electrode until it reaches the output. This is perfectly fine, but it makes it difficult to view the DNA in the electrode region. To illustrate the speed of the DNA migration, we first flow the DNA through the channel with the voltage off. Then, while viewing the fluorescence in the channel between the electrodes, the voltage is turned on, and rapid migration occurs toward the +V electrode.

7.4.2. Channel Electrodes for Isoelectric Focusing Combined with Field Flow Fractionation

Isoelectric focusing is an electrophoretic separation based on isoelectric point of proteins. The separation is done in a non-sieving medium (sucrose density gradient, agarose, or polyacrylamide gel) in the presence of carrier ampholytes, which establish a pH gradient increasing from the anode to the cathode during the electrophoreis. As the protein migrates into an acidic region of the gel, it will gain positive charge via protonation of the carboxylic and amino groups. At some point, the overall positive charge will cause the protein to migrate away from the anode (+) to a more basic region of the gel. As the protein enters a more basic environment, it will lose positive charge and gain negative charge, via ammonium and carboxylic acid group deprotonation, and consequently, will migrate away from the cathode (−). Eventually, the protein reaches a position in th pH gradient where its net charge is zero (defined as its pI or isoelectric point). At that point, the electrophoretic mobility is zero and is said to be focused.

Field-flow fractionation (FFF) is an elution chromatographic method for separating, concentrating, and collecting complex macromolecules, colloidal suspensions, emulsions, viruses, bacteria, cells, subcellular components, and surface-modified particles. In EFFF, an electric field, E, is applied across the channel and particles are subjected to the applied field according to their electrophoretic mobility. Fractionation occurs as different