1. IntroductionMicrofluidics refers to the behavior and control of liquids constrained to volumes near the micro liter range. In the late 1980s, which is the early stage of microfluidics, the development of microflow sensors, microflow sensors, micropumps and microvalves are dominating. With the introduction of life sciences and chemistry in the microfluidic field, which is proposed by Manz et al. at the 5th International Conference on Solid-State Sensors and Actuators, the field has been seriously developed. [1] Several competing terms, such as microfluidics, MEMSfluidics, or Bio-MEMS, and microfluidics appeared as the name for the new research discipline dealing with transport phenomena and fluid-based devices at microscopic length scales [2].

Figure 1.1: Size characteristics of microfluidic devices. [3]

According to Merriam-Webster, a fluid is a substance (as a liquid or gas) tending to flow or conform to the outline of its container. According to Merriam also, a liquid is a fluid that has no independent shape but has a definite volume and does not expand indefinitely and that is only slightly compressible, and gas is a fluid that has neither independent shape nor volume but tends to expand indefinitely. [3] Fluids are differs from solid in terms of response to the force. For solid, as long as the elastic limit is not exceeded, it will returns to its original state. However, the fluid subjected to shearing force will remain its new equilibrium position after the removal of shearing force. When the shear stress is directly proportional to the rate of strain within the fluid, the fluid is said to be Newtonian. The flowing of fluids can be explained by the properties of the fluid and the flow, basically split into 4 categories:1. Kinematic properties (Eg: angular velocity, acceleration and strain rate)1. Transport properties (Eg: viscosity, thermal conductivity, and diffusivity)1. Thermodynamics properties (Eg: pressure, temperature, and density)1. Others properties such as surface tension, vapor pressure, and surface accommodation coefficients. [3]

The study of fluid mechanics generally proceeds from the assumption that the fluid can be treated as a continuum. If the molecules are sparsely distributed relative to the length scale of the flow, assuming continuity of fluid and flow properties can is probably a dangerous approach. However, even at microscopic length scales, there can still be many thousands of molecules within a length scale significant to the flow. In order that a fluid can be modeled as continuum, all of its properties must be continuous. [4] The point quantities including the fluids kinematic properties, thermodynamic properties, transport quantities must also be continuous in order for the fluid to be treated as a continuum. For the transport quantities to behave continuously, it is important that the fluid molecules interact with themselves rather than with the flow boundaries. For gas flow, it is described by kinetic gas theory in which the gas molecule is considered to move in a straight line in constant speed until it collides with another molecule. There are some important parameters in assessing the state of fluid in motion, including Mach number (Ma), Knudsen number (Kn) and the Reynolds number (Re). The Mach number is a measure of the compressibility of a gas and can be thought of as the ratio of inertial forces to elastic forces. The Knudsen number has tremendous importance in gas dynamics which providing a measure of how rarefied a flow is, or how the density is relative to the length of flow. The Reynolds number is that it is a measure of the ratio between inertial forces and viscous forces in a particular flow.

Fluid drive or pumping methods include applied pressure drop, capillary pressure, electrophoresis, electro-osmosis, electrohydrodynamic force and magnetohydrodynamic force. For pressure drive, which is commonly used from the macroscopic world, is simply the application of positive pressure to one end of the channel or applied vacuum to the other end. Due to drag at the walls, the flow is slowest at the edges, causing a parabolic profile which having maximum velocity at the center.

Figure 1.2: Pressure driven flow. [5]

Another pumping force is wicking action of small diameter capillary force due to the surface tension. Capillary force is so commonly used and some people even never notice that the fountain of pen actually also using capillary action to control the flow of ink.

Figure 1.3: Even the pen we holding everyday are using the capillary to control ink flow. [6]

Electrophoretic flow can be induced only in liquids or gels with ionized particles. The application of a voltage across the ends of the channel produces an electric field along the channel that drives positive ions through the liquid toward the negative terminal and the negative ions to the positive terminal. Neutral particles are not affected by the field. The velocity of the ions is proportional to the field and charges, but inversely proportional to the size. [7] Velocity is inversely proportional to the viscosity in liquids, while for gels depends on the porosity.

Figure 1.4: Electrophoretic flow. [5]

Unlike pressure drive flow, the electro osmotic flow has almost flat velocity profile. Electro osmotic flow occurs because channels in glasses and plastics tend to have fixed charges on their surfaces. In glasses, the silanol (SiOH) groups at the walls are deprotonated in solution, leaving the surface with negative charges. [8] The negative ions then attract a diffuse layer of positive ions, forming a double layer in the liquid. The layer of positive ions is not tightly bound and can move under an applied electric field. When this sheath of ions moves, it drags the rest of the channel volume along with it, creating the electro-osmotic flow. This is benefits in many situations in biological analysis where the spreading of a short length sample into neighboring regions of channel is not desired. [5]

Figure 1.4: Electro-osmotic flow. [5]

2.0Microfluidic technology DNA Lab on a Chip (LOC)2.1Introduction of DNA Lab on a ChipLab-on-a-chio is becoming a revolutionary tool for many different applications in chemical and biological analyses due to its fascinating advantages (fast and low cost) over conventional chemical or biological laboratories. Besides, the simplicity of LOC systems will enable self-testing capability for patients or health consumers overcoming space limitation. The basic idea of LOC is to reduce biological or chemical laboratories to microscale system, hand-held size or smaller. With this microscale system, the reagents and chemicals used in biological and chemical reaction which are expensive can be minimized, and hence low cost. In addition, with the small amounts of sample, the reaction time can be faster. Another advantage is, only small amount of by products will be produced, this is very important for systems that produce harmful by products.

2.2 Design considerationTo design the DNA LOC, we should first understand the structure of DNA. The genetic code is stored in cell chromosomes, each containing long strands of deoxyribonucleic acid (DNA). [9-10] The building blocks of DNA are molecules called nucleotides that consist of a base joined to a sugar-phosphate backbone as shown in Figure 2.1. Each nucleotide molecule has two ends, labeled 3 and 5, corresponding to the hydroxyl and phosphate groups attached to the 3 and 5 positions of carbon atoms in the backbone sugar molecule as shown in Figure 2.2. In the long DNA chain, the 3 end of one nucleotide connects to the 5 end of the next nucleotide. [9-10]

Figure 2.1: Illustration of the twisted double-helix structure of DNA. [9-10]

Figure 2.2: The polymerase chain reaction (PCR). [9-10]

A primary objective of genetic diagnostics is to decipher the sequence of nucleotides in a DNA fragment after its extraction and purification from a cell nucleus. This task is difficult due to the miniscule concentration of DNA available from a single cell. As a solution, scientists resort to a special biochemical process called amplification to create a large number of identical copies of a single DNA fragment. The most common amplification method is the polymerase chain reaction (PCR). Invented in the 1980s by Kary Mullis, for which he was awarded the Nobel Prize in Chemistry in 1993, it allows the replication of a single DNA fragment using complementarity.The basic idea is to physically separatedenaturethe two strands of a double helix and then use each strand as a template to create a complementary replica. [5]

2.3 Device developmentPCR on a silicon chip was first demonstrated around 1994 by several groups, and by the end of the 1990s there had been several demonstrations of PCR on a chip. [11-12] This section describes silicon miniature PCR thermal cycling chambers developed at Lawrence Livermore National Laboratory (LLNL) of Livermore, California as in Figure 2.3.[13] Earlier designs had a polysilicon heater on a silicon nitride membrane for heating the fluid inside the chamber and used a separate, external temperature sensor. By changing the heater material to platinum, which is commonly used as a temperature sensor, both heating and sensing operations can be performed with the same platinum element. Testing of early devices showed that there were temperature variations as high as 10C across the chamber. By relocating the heater away from the membrane so that heat flows through the highly thermally conductive silicon walls of the chamber, the temperature uniformity of the fluid was greatly improved. A fan was added for more rapid cooling. These modifications have yielded much tighter closed-loop temperature control and enabled faster cycling, from around 35s per cycle to as little as 17s per cycle. These cycle times are far faster than the approximately 4 min per cycle needed in the industry-workhorse Applied Biosystems GeneAmp PCR System 9600 [14]. The LLNL system has detection capability in addition to amplification. In a variation of traditional PCR, the addition of TaqMan dyes (probes), which link to certain sections of a DNA strand (just like the primers), results in fluorescence of green light from each replicated DNA strand when excited by a blue or ultraviolet source [14, 15]. Thus, the intensity of the fluorescence is proportional to the number of replicated DNA strands matching the TaqMan probe in the solution. This procedure has the advantage of simultaneous DNA amplification and detection but only works when suitable primers and probe have been added to the solution for the type of DNA under test. Thus, the number of different DNA sections potentially being identified is equal to the number of PCR chambers that can be run simultaneously. In demonstrations at LLNL with different cells, there was no detectable fluorescence signal for the first 2025 cycles, depending on the initial concentration. After cycling on the order of 515 minutes, the signal appeared and rapidly grew if there was a match. [5]

Figure 2.3: The front side and the back side of an early micromachined silicon PCR chamber. A polysilicon heater on a silicon nitride membrane cycles the solution between the denaturing and incubation temperatures of PCR. [13]

2.4 Packaging and DesignPackaging of microfluidic chip is an important factor that completes the connection between the microfluidic chip and other systems such as fluidic, electrical, optical and etc (Figure 2.4). Since microfluidic chip interacts with fluidic samples, the packaging of microfluidic chips must also meet the fluidic flow requirements. In microfluidic packaging, polymers are generally used for encapsulation. Material compatibility based on biology and chemical stability based on the reagents is the key parameters of the package development [16]. Challenge in the microfluidic packaging is on the integration of different systems such as electrical, optical into a common platform and miniaturizing the system into a hand held system.

Figure 2.4: Schematic view of an integrated bio microfluidic package for DNA Lab On a Chip application. [17]

The system has been developed with integrated reservoir and valves for DNA (Deoxyribonucleic Acid) lab on a chip (LOC) application. A polymer material, PDMS (Polydimethylsiloxane) is used for encapsulating the DNA chip. Channels, reservoirs and valves are formed on the PDMS material by casting method. A mold fabricated in acrylic material is used. A double sided adhesive tape is used to bond PDMS substrates and DNA chip to PDMS substrate. [17] A capillary force passive valve has been designed and fabricated in PDMS. Threshold pressure required for valve opening has been characterized for different capillary channel dimensions. Reservoirs with different dimensions are fabricated based on the volume required in the DNA extraction process. Filled reservoirs with reagents are sealed by a PDMS membrane. An external actuation is used to apply pressure on the PDMS membrane to push the fluid from the reservoir to DNA chip. [17] Package requirements for a microfluidic chip are different from package for a microelectronic chip. In the microfluidic package, polymer material such as thermo-plastic or elastomer is used for formation of microfluidic components and chip encapsulation. Fluidic channels, reservoirs, valves are formed on the polymer substrate to transport fluid from the package inlet port to the fluidic chip (Figure 2.5).

Figure 2.5: Schematic view of microfluidic package with reservoirs and valves.[17]

The fluidic chip in this package is a DNA chip which has micro machined inlet and outlet holes. The chip consists of a filter, binder and a Polymer Chain Reaction (PCR) chamber. The DNA LOC is fabricated in silicon. The filter, binder and PCR chamber are formed by bulk micromachining method. The Silicon DNA LOC wafer is bonded to a glass wafer by anodic bonding method. The bonding process is conducted at the wafer level by using a wafer bonder. The glass wafer is used for optical approach detection. The filter is used to separate the blood cell which contains DNA from the blood sample and a binder is used to bind the DNA onto the chip. The bound DNA particles are removed from the chip by a process called elution. The collected elution sample containing DNA particles is mixed with the primer and injected into the PCR chamber of the chip for amplification and further analysis. [17]

Package has been designed with three PDMS substrate layers. Design of microfluidic components, like fluidic ports, valves and reservoirs, is based on the properties of PDMS material. Channels for fluidic flow are formed on the lower PDMS layer (Figure 2.6). Formation of channel on PDMS can be done by two methods, casting and soft lithographic method. The dimension of the fluidic channel is about 500 um. Casting method is used in making the substrate. An acrylic mold is designed and fabricated to obtain this substrate layer. Fluidic inlets and outlets to the micro channel have been designed on the upper substrate (Figure 2.6). The third substrate is used to form reservoir and valves. It is integrated on to the upper PDMS layer (Figure 2.7). Based on the DNA extraction protocol 4 reagents are required and hence 4 reservoirs are designed based on the reagent volume. Inlet ports are provided for blood sample and primer. These two fluidic samples are injected into the channel directly during the extraction process. [17]

Figure 2.6: The left figure showing lower PDMS substrate with channel and chip, while the right side showing upper PDMS substrate with inlet and outlet ports. [17]

Figure 2.7: cross section view of the system.

The complete package has reservoirs for reagents storage. A vertical channel at bottom of the reservoir links reservoir to package channel. A passive valve is embedded in the vertical channel. The valve is pressure activated. At storing condition, the valve is closed to prevent reagent flowing from reservoir to cartridge channel (Figure 2.8). Once fluidic pressure in reservoir increases and reaches the threshold pressure, the valve opens [17]. The advantage of the valve is passive and therefore controls the flow without any moving parts. The reservoir is covered by PDMS membrane. With external actuator piston, the membrane starts to deforming and therefore the fluidic pressure in reservoir increases. Once fluidic pressure reaches valves threshold pressure, fluid passes through the valve and flows to cartridge. The flow rate is controlled by adjusting pistons pushing speed.

Figure 2.8: Structure of a capillary force Passive valve.

2.5 IntegrationA microfluidic package has been developed with integrated reservoir and valves for a DNA lab on a chip application (Figure 2.9). There are 4 reservoirs have been formed on a PDMS substrate which has been filled with different reagents used for DNA extraction. A micro-valve is located between the reservoir and the channel in package. [17]

Figure 2.9: Microfluidic package with integrated reservoir and valve for DNA extraction. [17]2.6 FabricationThe entire microchip was developed on glass substrate. For this purpose, gold interdigitated microelectrodes for applying DC potential to the sample were fabricated over the glass by using photolithography and the evaporation method as shown in Figure 2.10. At first, photoresist AZ-1512 was spin-coated on glass and patterned using photolithography. After photolithography process, gold electrode was deposited using thermal evaporator. The electrode surface was cleaned with acetone and dried with gas. The microchannel was imprinted in the PDMS mold using the negative molding method. For this purpose, 40 um thick negative photoresist (SU-8 2075, Micro Chem.) was spin-coated and patterned on the silicon wafer. A degassed mixture of Sylgard 184 silicone elastomer along with curing agent (in 10:1 v/v ratio) was poured on the SU-8 patterned wafer and cured for 4 h at 72 C. The PDMS mold formed was then peeled off manually and drilled to produce access holes of 3 mm diameter. The width and depth of the microchannel in the entire chip were 250 um and 200 um, respectively. [18]

Figure 2.10: Schematics for fabrication of cell lysis and PCR modules on glass substrate using the photolithographic technique. [18]

For fabrication of PCR module, the PDMS microchannel was fabricated once again by the negative molding method. SU-8 was first spin-coated onto a bare silicon wafer and patterned to make microchannel using photolithography with mask aligner (MA-6, Karl-Suss) (Figure 2.10). The degassed PDMS monomer mixture and curing agent (10:1 v/v) was poured on the SU-8 negative master pattern and cured for 4 h at 72 oC. The PDMS was then peeled off and manually drilled to produce access holes. The width and depth of the microchannel were once again 250 and 200 um, respectively, and the total length of microchannel was 1550 mm including cell lysis module (114 mm) and for 20 PCR cycles. The ratio of the channel lengths of the three different temperature zones for thermocycling, namely denaturation (92 oC), annealing (55 oC, and extension (73 oC was 2: 2: 3. It ensured a retention time of 30 s in denaturation and annealing zones and 45 s in the extension zone for PCR premix flowing in the microchannel at 5 ul/min rate. The PCR module was finally interconnected with the cell lysis module using silicone tubes inserted into drilled access holes. The ITO heater electrode was the material of choice for thermal cycling due to its property showing linear variation of its temperature by application of DC power and was fabricated using conventional photolithography and wet etch process [17]. For this purpose, positive photoresist (AZ1512, Clariant) was spin-coated on glass with deposited ITO film (Samsung Corning) and then photoresist AZ1512 was patterned using photolithography to make ITO electrodes. The ITO film was then etched using FeCl3/HCl solution for 2 h and photoresist was removed. For electrical isolation, a thin layer of PDMS was spin coated over microheater surface and baked at 95oC for 30 min. ITO heaters were calibrated for liquid/air temperature control by inserting thermocouple into the microchannel during UV-ozone bonding. For finalizing the device fabrication steps, the PDMS mold and glass substrate containing ITO/Gold electrodes were bonded with each other by UV-ozone treatment for 40 min. Fig. 1 shows the schematics of fabrication process of continuous-flow PCR chip, while Fig. 2 illustrates the fabricated microchip. [18]

3.0References[1]Manz, A., Graber, N., and Widmer, H. M., Miniaturized Total Chemical Analysis Systems: A Novel Concept for Chemical Sensing, Sensors and Actuators B, Vol. 1, 1990, pp. 244-248.

[2] Gravesen, P., Branebjerg, J., and Jensen, O.S., Microfluidics A Review, Journal of Micromechanics and Microengineering, Vol. 3, 1993, pp. 168-182.

[3]Nam-Trung Nguyen,Steve Wereley,Fundamentals and Applications of Microfluidics, Artech House Microelectromechanical Systems Library, 2002.

[4] Deen, W. M., An Analysis of Transport Phenomena, Oxford, UK: Oxford University Press, 1998.

[5]Nadim Maluf, An Introduction to Microelectromechanical Systems Engineering, Artech House Microelectromechanical Systems Library, 2004.

[6]Prof. Dr. Roland Zengerle and Stefan Haeberle, Introduction to microfluidic, slide 33.

[7]Kovacs, G. T. A., Micromachined Transducer Sourcebook, Boston, MA: WCB McGraw-Hill, 1998, Section 6.6.

[8]Sharp, K. V., et al., Liquid Flow in Microchannels, in The MEMS Handbook, M. Gad-el-Hak (ed.), Boca Raton, FL: CRC Press, 2002, Chapter 6.

[9]Stryer, L., Biochemistry, New York: W. H. Freeman and Co., 1988, pp. 7190,120123.

[10]Darnell, J., L. Harvey, and D. Baltimore, Molecular Cell Biology, 2nd ed., New York: Scientific American Books, 1990, p. 219.

[11]Yao, J. J., et al., A Low Power/Low Voltage Electrostatic Actuator for RF MEMS Applications, Technical Digest of Solid-State Sensor and Actuator Workshop, Hilton Head, NC, June 2000, pp. 246249.

[12]Yao, J. J., RF MEMS from a Device Perspective, Journal of Micromechanics and Microengineering, Vol. 10, 2000, pp. R9R38.

[13]Ulrich, R., and L. Schaper, Putting Passives in Their Place, IEEE Spectrum, Vol. 40, No. 7, July 2003, pp. 2630.

[14]Van Schuylenbergh, K., et al., Low-Noise Monolithic Oscillator with an Integrated Three-Dimensional Inductor, Technical Digest of International Solid-State Circuits Conference, San Francisco, CA, February 2003, pp. 392393.[15]Chua, C. L., et al., Out-Of-Plane High-Q Inductors On Low-Resistance Silicon, Journal of Microelectromechanical Systems, Vol. 12, No. 6, December 2003, pp. 989995.

[16]S C Chong, Ling Xie, et al, Disposable Polydimethylsioxane Package for Bio-Microfluidic System, Electronic Components and Technology Conference, May 2005, pp: 617-621.

[17]Ling Xie, et al.,Development of an Integrated Bio-Microfluidic Package with Micro-Valves and Reservoirs for a DNA Lab on a Chip (LOC) Application, Electronic Components and Technology Conference,2006.

[18]Sandeep Kumar Jha, et al., Development of PCR Microchip for Early Cancer Risk Prediction IEEE SENSORS JOURNAL, VOL. 11, NO. 9, SEPTEMBER 2011