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420MICROFLUIDICS

207.Pettersson A, et al. A feasibility study of solid supported enhanced microdialysis. Analy Chem 2004;76:1678–1682.

208.Ao X, Sellati TJ, Stenken JA. Enhanced microdialysis relative recovery of inflammatory cytokines using antibodycoated microspheres analyzed by flow cytometry. Anal Chem 2004;76:3777–3784.

209.Kissinger CB, Kissinger PT. Can preclinical ADMET-PK now be done more efficiently and effectively? Preclinica 2004;2: 319–323.

210.Olson DL, Lacey ME, Sweedler JV. High-resolution microcoil NMR for analysis of mass-limited, nanoliter samples. Anal Chem 1998;70:645–650.

211.Wolters AM, Jayawickrama DA, Sweedler JV. Microscale NMR. Curr Opin Chem Biol 2002;6:711–716.

212.Khandelwal P, et al. Studying rat brain Neurochem using nanoprobe NMR spectroscopy: A metabonomics approach. Analy Chem 2004;76:4123–4127.

Further Reading

The most comprehensive sources for microdialysis sampling are the book and the two separate journal issues shown below.

Robinson T, Justice JB, editors. Microdialysis in the Neurosciences. Amsterdam (The Nethernand): Elsevier; 1991.

Lunte CE Anal Chim Acta 1999;379:227–369.

Elmquist WF, Sawchuk RJ. Microdialysis sampling in drug delivery. Adv Drug Del Res 2000;45:123–307.

See also ELECTROPHORESIS; GLUCOSE SENSORS; HYDROCEPHALUS, TOOLS FOR DIAGNOSIS AND TREATMENT OF; PHARMACOKINETICS AND PHARMACODYNAMICS.

MICROFLUIDICS

GLENN M. WALKER

North Carolina State University

Raleigh, North Carolina

INTRODUCTION

Microfluidics is the study and application of fluids at the microscale. The most common definition of the microscale is that one or more device dimension be in the range of 1–1000 mm. For reference, the diameter of an average human head hair is 150 mm, the average thickness of a human fingernail is 360 mm, and the diameter of a human red blood cell is 7 mm. Miniaturization technology, originally developed by the microelectronics industry, has been used to create microscale fluid components and complete microfluidic systems with pumps, valves, and filters, incorporated onto single microchips have been demonstrated.

By applying the analogy of the microelectronics industry (i.e., continuously incorporating more features into smaller areas) a logical application of microfluidics is to create lab-on-a-chip (LOC) systems. Lab-on-a-chip systems, also known as micro-total-analysis systems (mTAS), incorporate the functionality of biology or chemistry laboratories onto a single microfabricated chip. Ideally, a LOC system would be able to execute all of the tasks routinely performed in a biology or chemistry laboratory, such as sample preconditioning, mixing, reaction, separa-

tion, and analysis. Laborand time-intensive procedures would be reduced to instant results derived from a series of automated steps performed on a LOC.

Microscale fluid handling confers many advantages over traditional lab operations (1). First, fluid quantities ranging from picoliters to microliters are used, thus reducing the amount of sample required for tests. Second, the amount of time required to perform some analyses (e.g., capillary electrophoresis) is reduced to seconds, which means analyses can be conducted many times faster than with traditional methods. Third, devices can be manufactured using microfabrication technology, which translates into reduced cost per device; disposable LOC systems can easily be envisioned.

In general, microfluidic devices are in early stages of development and are most often found in academic research laboratories. However, the benefits of these systems have been exploited to develop new medical devices for clinical diagnostics and point-of-care testing. Commercial examples of devices that make use of LOC concepts are discussed at the end of this article.

THEORY

Fluid Mechanics

The term microfluidics encompasses both liquid and gas behavior at the microscale, even though in most applications the working fluid is a liquid. All of the concepts discussed here are directed toward liquids. Other works are available which provide information on gas behavior at the microscale (2).

Fluid behavior at the microscale is different from that commonly observed in everyday experiences at the macroscale, owing primarily to the very low Reynolds (Re) numbers of the flow regime plus the large surface area/volume (SAV) ratios of the flow domain. As a consequence, viscous forces and surface tension effects become dominant over fluid inertia, and transport phenomena are purely diffusive.

Fluid flow at the microscale is typically laminar. Fluid flows are classified based on their flow regime, which can be predicted with the Re number. The Re number is the ratio of inertial forces to viscous forces and can be calculated with the equation

Re ¼

rVDh

(1)

m

where r is the fluid density, V is the characteristic fluid velocity, Dh is the hydraulic diameter of the microchannel, and m is the fluid viscosity. Fully developed fluid flow in a channel of circular cross-section is considered laminar if the Re number is <2100. For Re numbers between 2100 and 2300, the flow is considered transitional: it shows signs of both laminar and turbulent flow. A Re number >2300 indicates turbulent flow.

Laminar flow is predictable in the sense that the trajectories of microscopic particles suspended in it can be accurately predicted (Fig. 1a). Particles suspended in a turbulent fluid flow behave chaotically and their position as a function of time cannot be accurately predicted

MICROFLUIDICS 421

Figure 1. (a) Particles suspended in a laminar flow within a straight microchannel follow straight trajectories. (b) Particles suspended in a turbulent flow within a straight microchannel do not follow straight trajectories unless they are very close to the wall.

(Fig. 1b). Fluid flow that has a Re number <1 is also known as viscous flow, creeping flow, boundary-layer flow, or Stokes flow.

Low Re number flow is best visualized by imagining how honey (or any viscous substance) behaves when poured or stirred. For example, water flowing in microchannels will generally have a Re number <1. In this case, water will behave like a very viscous liquid (i.e., like honey). An important point to make here is that the properties of water do not change at the microscale; rather the microscale dimensions involved make the water appear more viscous than what we are accustomed to at the macroscale. An excellent description of low Re number environments has been given by Purcell (3). Very viscous fluid flows have certain characteristics: the flow is reversible, mixing is difficult, and flow separation does not occur

(4).

Reversibility is the ability of a suspended particle in a fluid to retrace its path if the flow is reversed. This is a result of the minimal inertia (i.e., low Re number) present in fluid flows at the microscale. Figure 2 shows the path a suspended microscopic particle might take in forward and reverse flow.

Figure 3. (a) At low velocities (low Re numbers), flow separation will not occur in microchannels. (b) At higher velocities (larger Re numbers), flow separation may become apparent.

A second characteristic of microscale fluid flow is a lack of flow separation. Flow separation is commonly observed in the form of vortices, which are recirculating flows separate from the main flow. Because of the low Re number environment, vortices usually will not form within microfluidic channels, as shown in Fig. 3a. Separation will only occur in flows wherein inertial forces are significant relative to viscous forces (Re > 1). Figure 3b is a qualitative sketch of flow separation in a cavity.

The third characteristic of microscale fluid flows is inefficient mixing as a result of very low Re number flow, and thus negligible inertia. Low inertia means that stirring is not effective and that mixing must be accomplished by diffusion. At the macroscale, stirring minimizes the diffusion distances between two or more liquids by distributing ‘‘folds’’ of the liquids throughout the volume. Microscale methods of mixing have been developed that take advantage of the unique properties of the scale and improve the efficiency of mixing over simple diffusion; examples include using three-dimensional (3D) channel geometries, patterned channel surfaces, and pulsatile flow (5).

Figure 4 shows two streams flowing down a microchannel side-by-side. Because of the low Re number environment the streams will only mix by diffusion. If the flowrate is slow enough, the streams will eventually become uniformly mixed across the whole microchannel width.

The hydraulic diameter, Dh, of a microchannel is determined by its cross-sectional geometry and can be calculated with the equation

Dh ¼

4A

(2)

P

Figure 2. (a) A suspended particle in laminar flow around an obstacle in a microchannel. (b) If the flow is reversed, the particle will retrace the same path.

Figure 4. Two streams flowing in a microchannel will only mix by diffusion. Note that the concentration across the width of each half of the main microchannel is not constant as diffusional mixing progresses.

422 MICROFLUIDICS

Table 1. Diffusion Coefficients for Biologically Important Molecules in Watera

Molecule

T, 8C

D, cm2 s 1

Diffusion time, s

Cl

25

2.03 10 5

0.02

O2

18

2 10 5

0.03

Kþ

25

1.96 10 5

0.03

Naþ

25

1.33 10 5

0.04

Glucose

20

6 10 6

0.08

Lactose

20

4.3 10 6

0.12

Insulin

20

1.5 10 6

0.33

Hemoglobin

20

6.3 10 7

0.79

Urease

20

3.4 10 7

1.47

aAll values are from Ref. 6. The time for each particle to diffuse 10 mm is shown for comparison.

where A and P are the microchannel cross-sectional area and wetted perimeter, respectively. The hydraulic diameter is often used to calculate important flow characteristics for noncircular microchannel cross-sections.

At microscale dimensions diffusion is an effective mechanism for transporting molecules because of the relatively short distances involved. Particles diffuse from areas of high concentration to areas of low concentration and will eventually diffuse to uniform concentration throughout a given volume. The mean distance, d, a particle travels in a time, t, can be predicted with the equation

d2 ¼ 2Dt

(3)

where D is the diffusion coefficient of the particle. Diffusion times are proportional to the square of distance, which means that particles can diffuse across microscale distances within a particular medium in a matter of seconds. Table 1 lists representative molecules of biological significance and their diffusion coefficients.

The SAV ratios become very large at the microscale. Typical SAV ratios for macroscale containers such as Petri dishes or culture flasks are 10 cm 1, while they are 800 cm 1 for microfluidic channels. Increased SAV ratios allow diffusion-limited processes, such as immunoassays to become much more efficient at the microscale because of the increased surface area available for binding. Large SAV ratios also allow rapid heat radiation from microscale fluid volumes and efficient gas exchange with the ambient atmosphere and fluid in microchannels (assuming the microchannel is made of a gas-permeable material). Enhanced gas transport is a critical ingredient for cell culture in microscale environments. One drawback of large SAV ratios is that evaporation becomes a significant problem.

The surface tension of a liquid becomes increasingly important at very small dimensions. To visualize this, think of a liquid surface as an elastic skin. If a slit were made in that skin, a certain amount of force per unit length would be required to hold the two sides of the slit together. The amount of force required to hold the two sides together is called the surface tension. Because the liquid surface is under tension, liquid confined by the surface (e.g., a rain-

drop) will experience an internal pressure. This pressure is called the Young–LaPlace pressure. Smaller fluid volumes result in larger SAV ratios, thus increasing the internal pressure. The pressure within a drop of liquid can be calculated with the formula

DP ¼

2g

(4)

R

where g is the liquid surface energy and R is the radius of the drop. At microscale dimensions, significant pressures can be created by surface tension. A common result of the pressure difference of an air/liquid interface is the capillary effect: a pressure difference across the interface propels liquid through a small diameter capillary or microchannel.

The capillary effect also depends on the contact angle of the microchannel surface. Hydrophobic surfaces (e.g., polymers) have contact angles >908 and hydrophilic surfaces (e.g., glass) have contact angles <908. Microfluidic devices with hydrophilic surfaces can be filled via capillary action. The pressure difference at an air–liquid interface within a microchannel with square cross-sectional area can be cal-

culated with the formula

 

 

 

 

DP

¼

2g

cosðuc

Þ þ

cosðuc

(5)

 

 

W

H

Þ

where W and H are the microchannel width and height, respectively, and uc is the contact angle of the liquid on the internal microchannel walls. Conversely, equation 5 gives the pressure required to force water into a hydrophobic microchannel of rectangular cross-section.

Microfluidic Modeling

Microscale fluid flow can be modeled from either a macroscopic or microscopic vantage point. Macroscopic modeling treats the fluid as a well-mixed volume while the microscopic view looks at how particles suspended in the fluid would behave under different flow conditions.

Macroscopic modeling, also called lumped modeling, uses conservation of mass to predict microfluidic system behavior. A pressure drop, DP, applied across a microchannel (or other conduit) with fluidic resistance Z, will induce a volumetric flow rate Q:

DP ¼ QZ

(6)

All microchannels have a fluidic resistance associated with them that depends on the geometry of the microchannel and the viscosity of the fluid. The fluidic resistance of a microchannel with a circular cross-section is given by

8mL

 

Z ¼ pR4

(7)

where m is the fluid viscosity, L is the microchannel length, and R is the microchannel radius. The fluidic resistance of a microchannel with a rectangular crosssection is given by

 

4mL

 

a

1

Z ¼

 

f

 

 

(8)

ab3

b

where f(a/b) is calculated with the formula

 

 

a

 

16

 

 

X

 

 

 

1024b 1 tanh ma

 

f

 

¼

 

 

 

 

 

(9)

b

3

ap5 n¼0 ð2n þ 1Þ5

When calculating the resistance of microchannels with rectangular cross-section, m is the fluid viscosity, L is the microchannel length, 2a and 2b are the microchannel width and height, respectively, and m is calculated with

m

¼

 

2n þ 1Þ

(10)

2b

 

 

 

If the aspect ratio of the microchannel is very small (i.e., 2b 2a) then the simplified formula

Z ¼

3mL

(11)

4ab3

can be used. The general rule of thumb is that equation 11 should be used for microchannels with b/a <0.1. The resistance of other geometries can be found elsewhere (7).

In predicting microfluidic system behavior, the analogies to Kirchhoff’s laws are used. The sum of pressure drops in a fluidic loop must be equal to zero; the total volumetric flowrate entering a node must be equal to the total volumetric flowrate leaving a node.

In contrast to the macroscopic view, microscopic modeling allows fluid behavior to be predicted. Specifically, the microscopic view allows the velocity profiles of a fluid flow to be calculated. Velocity profiles are plots that show the relative velocities of different portions of a fluid within a microchannel. Figure 1a is an example of a velocity profile.

The velocity of flow in a microchannel with circular cross-section varies radially and can be predicted with the formula

vðrÞ ¼

R2DP

1

r2

 

(12)

4mL

R2

where m is the fluid viscosity, L is the microchannel length, DP is the pressure drop, and R is the microchannel radius. The velocity profile of flow in a microchannel with rectangular cross-section varies along the height and width axes and can be predicted with the formula

vðx; yÞ ¼ 2mL

b2 y2 b

1

ð 1Þn m3 cosh ma

!

 

DP

4

X

 

1 cos my cosh mx

 

 

 

 

n¼0

 

 

(13)

 

 

 

 

 

 

where m is the fluid viscosity, L is the microchannel length, DP is the pressure drop, m is calculated from equation 10, and 2a and 2b are the microchannel width and height, respectively.

MICROFLUIDICS 423

Microscopic modeling is performed when precise modeling of fluid behavior is needed. For example, cells attached to the wall of a microchannel might affect flow; modeling at the microscopic level would reveal any perturbations of the flow caused by the cell. In contrast, macroscopic modeling is performed when the behavior of the entire microfluidic system is needed. For example, fluid flow in many parallel microchannels might be required. Macroscopic modeling would reveal the relative flowrates through each microchannel and provide the microchannel dimensions needed to guarantee equal flow through each.

PUMPING FLUIDS

Fluids are pumped through microfluidic channels by creating gradients; the two most common types being pressure and electrical. Other types of gradients and their applications are discussed elsewhere (8).

Pressure gradients are the most common method used to pump fluid. Pressure is applied to one end of a microchannel which causes the fluid to flow down the pressure gradient. Common methods for creating a pressure gradient include pumps or gravity. Most methods for creating pressure-driven flow use macroscale pumps attached to the microfluidic device via tubing. Ideally, pumps should be incorporated on-chip to realize the ultimate vision for LOC devices. Many types of microfluidic pumps have been demonstrated and they presently constitute an active area of research (9).

Pressure-driven flow is attractive for use in microfluidics because it is easy to set up and model. Some drawbacks for using pressure-driven flow are sensitivity to bubbles, sensitivity to motion (via the tubing connecting pumps to the microfluidic device), and parabolic flow profiles. Shear stress is proportional to the pressure drop across a microchannel, which should be taken into account when manipulating cells.

The other common way to pump fluids is to use electrical gradients. This method of pumping fluid is only practical at the microscale level because of the large electric fields and SAV ratios required. Pumping via electric fields is called electrokinetic flow and is based on two phenomena: electrophoresis and electroosmosis. Electrophoresis operates on the principle that charged particles in an electric field will feel a force proportional to the field strength and their charge. The particles will move through the electric field toward the pole of opposite charge. Larger particles move slower than smaller particles because of the drag produced by moving through a fluid. Figure 5a shows an example of electrophoretic flow.

A charged particle in an electric field of strength E will travel with a velocity equal to

v ¼ mepE

(14)

where mep is the electrophoretic mobility. Electrophoretic velocities are typically much smaller than the velocities caused by electroosmosis.

Electroosmosis will only function in the presence of an electric double layer at the surface of the microchannel.

424 MICROFLUIDICS

Figure 5. (a) Electrophoresis. Charged particles will move toward oppositely charged poles in an electric field. (b) Electroosmosis. Charges lining a microchannel surface will move with an applied electric field, thus inducing bulk flow via momentum transfer within the liquid. (c) An electric double layer forms at charged microchannel surfaces; the layer thickness is called the Debye length.

An electric double layer forms at a charged surface when oppositely charged particles from an electrically neutral liquid gather at the surface. The thickness of the electric double layer is known as the Debye length; the concentration of charged particles at a surface falls off rapidly as a function of distance and is shown in Fig. 5c. When an electric field is applied across the length of the microchannel, the ions gathered at the microchannel surface begin to slide toward the oppositely charged pole, as shown in Fig. 5b. As the ions slide, they drag their neighbors within the bulk liquid, toward the middle of the microchannel. The friction between subsequent sliding layers of ions causes the bulk fluid to begin moving.

If the Debye length is much less than the characteristic dimensions of the microchannel, then the bulk fluid velocity can be predicted with the equation

v ¼ meoE

(15)

where E is the electric field strength and meo is calculated with

meo ¼

ez

(16)

4pm

where e is the dielectric constant of the fluid, z is the zeta potential of the surface, and m is the fluid viscosity.

Electrokinetic flow is attractive because it only requires the integration of electrodes in a microfluidic device, which is straightforward by microfabrication standards. Electrokinetic flow is therefore amenable to interfacing with electronic control circuitry. Electrokinetic flow also results in a blunt flow profile, which reduces the distortion of transported samples. Lastly, electrokinetic flow has

very rapid response times, since the electrodes are integrated on chip, and are generally insensitive to movement off chip. Drawbacks to electrokinetic flow include fouling of the electrodes, which reduces electric field strength and therefore flowrate. Also, protein adsorption to microchannel surfaces affects the Debye layer, and in turn flow. Unintended side effects from electric fields on biological cells within the microfluidic device may also exist. Figure 6 shows the difference between the parabolic flow profile of pressure-driven flow and the blunt profile of electrokinetic flow.

Other methods of fluid flow based on surface tension, heat, and evaporation, have also been demonstrated. However, these methods have yet to be widely adopted and it is

Figure 6. (a) Pressure-driven flow is parabolic; the middle of the stream moves faster than regions near the wall. (b) Electrokinetic flow is blunt; all parts of the stream move at equal velocity. Note that residual pressures can cause the profile to become slightly parabolic in the direction of the negative pressure gradient.

unclear if they will prove more attractive than pressuredriven or electrokinetic flow.

FABRICATION

Fabricating microfluidic channels in traditional microfabrication materials, such as silicon and glass, can be achieved in two different ways. In the first, material is selectively removed, or etched, from a bulk substrate. The etched substrate is then bonded to another material (e.g., glass or silicon), which may have access holes or other features embedded in it. The result is an enclosed channel structure, as shown in Fig. 7a. The second method is to selectively add material to a substrate, and then bond another substrate to it. This method will also form enclosed channel structures as shown in Fig. 7b.

Photolithography is a fundamental part of all microfabrication (Fig. 8). Light is used to project patterns onto a photosensitive chemical, called a photoresist. The photoresist can be either positive or negative. Light chemically alters positive photoresist and makes it soluble in a developer. Negative photoresist is cross-linked by light, which makes it insoluble in developer. The patterned photoresist can be used as an etch mask for substrates, producing microchannels like those shown in Fig. 7a. Patterned photoresist can also be used in subsequent steps to direct the patterning of other materials that

MICROFLUIDICS 425

Figure 7. (a) Material is etched from a substrate and an enclosed structure is formed by bonding the etched substrate to a glass lid. (b) Walls of microchannels are built on top of a substrate and a glass lid is placed on top of the photoresist to form an enclosed structure.

cannot be directly patterned with light. In-depth treatments of microfabrication techniques can be found elsewhere (10,11).

Polymers have recently become popular alternatives to traditional (e.g., silicon or glass) microfabrication materials. Polymers can be used to make microchannels by using the same methods mentioned previously for silicon and glass. Polymers also have the advantage that they can be molded that makes them a cheaper alternative (relative to silicon or glass) for mass production.

Polymer microfluidic devices can be created by molding, hot embossing, injection molding, photopolymerization, and laser ablation or laser cutting. An attractive method for fabricating polymer microfluidic devices is to use a process known as micromolding (12). In this process, photolithography is used to make a pattern of the microchannels (called a master). The photoresist provides a positive relief from which a polymer mold can be cast. The polymer is poured over the master and allowed to cure. The polymer mold is then peeled from the master and either placed on a substrate or incorporated into a multilayer device. The two advantages of this microfabrication method are (1) no special microfabrication equipment is required, and (2) many inexpensive copies of a microfluidic device can be rapidly manufactured.

Drawbacks to using polymers include leaching of material into microfluidic channels, solvent incompatibility,

Figure 8. Photolithography requires an ultraviolet (UV) light source, a mask, and a photosensitive layer of material (i.e., photoresist). The photoresist is patterned with the UV light via a mask.

426 MICROFLUIDICS

Figure 9. (a) The Bioanalyzer uses microfluidic chips etched from glass. (Images courtesy of Agilent Technologies). (b) The etched glass chips are encased in a plastic assembly that facilitates handling and limits contamination. Images courtesy of Agilent Technologies. (c) Samples are loaded onto the chip and the chip is loaded into the Bioanalyzer. (Images courtesy of Agilent Technologies). (d) The Bioanalyzer then performs an analysis on the samples. (Images courtesy of Agilent Technologies.)

and the ability of some substances to diffuse into the polymer. Also, surface treatments are occasionally required to make polymers compatible with electrokinetic flow; silicon or glass have an inherent surface charge that allows them to be used in electrokinetic flow applications.

BIOMEDICAL APPLICATIONS OF MICROFLUIDICS

Microfluidic concepts have already been incorporated in a variety of biomedical devices (13). One example of a microfluidic device now found in many biomedical labs is Agilent’s Bioanalyzer. The Bioanalyzer system uses disposable chips etched in glass. The samples to be separated and reagents for the separation are loaded onto the chip via reservoirs. The chip is then placed in a reader, where electrokinetic flow is used to manipulate the samples in the reservoirs (Fig. 9).

Microfluidic capillary electrophoresis systems are becoming commonplace in laboratories. By using very small volumes for separation, joule heating from electrophoresis is rapidly radiated away from the gel.

Efficient heat radiation allows larger voltages to be used which translates into faster separations. The shorter separation distances used also contribute to reduced analysis times. Microfluidic capillary electrophoresis systems allow DNA to be rapidly analyzed, and have highly reproducible results since the entire process is automated. Lastly, contamination is minimized because the devices are disposable.

Because of their small size, another attractive aspect of capillary electrophoresis (CE) systems is that they can be incorporated into LOC devices and made part of a complete system. An example that has been demonstrated in several research labs is a system that takes cells as inputs, lyses them, performs all necessary preprocessing, DNA amplification, and so on, and then performs the DNA separations, all with no human intervention (14).

Microfluidics are also being used in clinical devices; devices for hematology and disposable assays for point- of-care diagnostics are among those now being researched and brought to market. A handheld point-of-care device made by i-STAT is an example of a clinical microfluidic device (Fig. 10). The handheld device quantifies analytes in

(a)

Figure 10. (a) The handheld point- of-care device manufactured by i- STAT performs analyses on blood samples contained in a disposable cartridge. Reprinted with permission from ACS (from Ref 15). Copyright 1998 American Chemical Sa. (b) Microfluidic cartridges are loaded with a sample and then plugged into the i-STAT handheld device. Image courtesy of Abbott Point-of-Care. The cartridge contains all necessary microfluidic control and sensing components that are then actuated by the handheld device.

Cartridge label

Sample entry well gasket

Fluid channel

Cartridge cover

Sample entry well

Tape gasket

Biosensor chips

Calibrant pouch

Puncturing barb Cartridge base Air bladder

(b)

MICROFLUIDICS 427

Cartridge

blood samples that have been deposited on a disposable chip. The general procedure for operation is given below (15).

A patient’s blood sample is deposited in a well on the disposable chip and then the well gasket is snapped shut (Fig. 10a). The microfluidic chip is then inserted into a handheld reader that performs an automated analysis of the blood sample, as shown in Fig. 10b. On-chip biosensors are automatically calibrated and checked for accuracy with an on-chip packet of calibrated solution. Once their accuracy has been determined, the calibration solution is flowed to the waste compartment. The blood sample is then flowed over the biosensors and the concentrations of different analytes are displayed on the handheld device screen within a few minutes. Diaphragm pumps are used to move the fluid. A variety of chips are available for different assays, including electrolytes and blood gases.

CONCLUSION

Microfluidics is the study and application of fluids at the microscale. Techniques used by the microelectronics industry have been adapted to facilitate the creation of micron-size channels capable of carrying fluid. The physical behavior of fluid at the microscale differs from behavior observed at the macroscale in everyday experience. The miniaturization of fluid handling has allowed LOC devices to be created in which all of the procedures of a traditional chemistry or biology lab are performed automatically in a single microfabricated chip. Lab-on-a-chip devices will allow new clinical and research tools to be developed.

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4.Meldrum DR, Holl MR. Tech.Sight. Microfluidics. Microscale bioanalytical systems. Science 2002;297(5584):1197–1198.

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6.Stein WD. Channels, Carriers, and Pumps: An Introduction to Membrane Transport. San Diego: Academic; 1990.

7.Shah RK, London AL. Laminar Flow Forced Convection in Ducts. New York: Academic; 1978.

8.Stone HA, Stroock AD, Ajdari A. Engineering flows in small devices: Microfluidics toward a lab-on-a-chip. Ann Rev Fluid Mech 2004;36:381–411.

9.Laser DJ, Santiago JG. A review of micropumps. J Micromech Microeng 2004;14(6):R35–R64.

10.Kovacs G. Micromachined Transducers Sourcebook. Boston: WCB McGraw-Hill; 1998.

11.Madou M. Fundamentals of Microfabrication. 2nd ed Boca Raton(FL): CRC Press; 2002.

12.McDonald J, et al. Fabrication of microfluidic systems in poly(dimethylsiloxane). Electrophoresis 2000;21(1):27–40.

13.Beebe DJ, Mensing GA, Walker GM. Physics and applications of microfluidics in biology. Ann Rev Biomed Eng 2002;4:261–286.

14.Waters L, et al. Microchip device for cell lysis, multiplex PCR, amplification, and electrophoretic sizing. Anal Chem 1998; 70(1):158–162.

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See also DRUG DELIVERY SYSTEMS; NANOPARTICLES