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Biotechnology Advances 25 (2007) 483 ­ 514 www.elsevier.com/locate/biotechadv

Research review paper

Micropumps, microvalves, and micromixers within PCR microfluidic chips: Advances and trends

Chunsun Zhang, Da Xing , Yuyuan Li

MOE Key Laboratory of Laser Life Science & Institute of Laser Life Science, South China Normal University, No.55, Zhongshan Avenue West, Tianhe District, Guangzhou 510631, PR China Received 26 March 2007; received in revised form 6 May 2007; accepted 17 May 2007 Available online 23 May 2007

Abstract This review surveys the advances of microvalves, micropumps, and micromixers within PCR microfluidic chips over the past ten years. First, the types of microvalves in PCR chips are discussed, including active and passive microvalves. The active microvalves are subdivided into mechanical (thermopneumatic and shape memory alloy), non-mechanical (hydrogel, sol­gel, paraffin, and ice), and external (modular built-in, pneumatic, and non-pneumatic) microvalves. The passive microvalves also include mechanical (in-line polymerized gel and passive plug) and non-mechanical (hydrophobic) microvalves. The review then discusses mechanical (piezoelectric, pneumatic, and thermopneumatic) and non-mechanical (electrokinetic, magnetohydrodynamic, electrochemical, acoustic-wave, surface tension and capillary, and ferrofluidic magnetic) micropumps in PCR chips. Next, different micromixers within PCR chips are presented, including passive (Y/T-type flow, recirculation flow, and drop) and active (electrokinetically-driven, acoustically-driven, magnetohydrodynamical-driven, microvalves/pumps) micromixers. Finally, general discussions on microvalves, micropumps, and micromixers for PCR chips are given. The microvalve/micropump/micromixers allow high levels of PCR chip integration and analytical throughput. © 2007 Elsevier Inc. All rights reserved.

Keywords: Polymerase chain reaction (PCR); Microfluidic chip; Microvalves; Micropumps; Micromixers

Contents 1. 2. Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Fluid switching: microvalves for PCR microfuidic chips . . . . . . . . . . . . . . . . . 2.1. Active mechanical microvalves -- thermally actuated microvalves for PCR chips 2.1.1. Thermopneumatic microvalves . . . . . . . . . . . . . . . . . . . . . . 2.1.2. Shape memory alloy microvalves . . . . . . . . . . . . . . . . . . . . . 2.2. Active non-mechanical microvalves -- phase change microvalves for PCR chips 2.2.1. Hydrogel microvalves . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.2.2. Sol­gel microvalves . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 484 485 485 485 485 486 486 487

Corresponding author. Tel.: +86 20 85210089; fax: +86 20 85216052. E-mail address: [email protected] (D. Xing). 0734-9750/$ - see front matter © 2007 Elsevier Inc. All rights reserved. doi:10.1016/j.biotechadv.2007.05.003

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2.2.3. Paraffin microvalves . . . . . . . . . . . . . . . . . . . 2.2.4. Ice microvalves . . . . . . . . . . . . . . . . . . . . . 2.3. Active external microvalves for PCR chips . . . . . . . . . . . . 2.3.1. Modular built-in microvalves . . . . . . . . . . . . . . 2.3.2. Pneumatic microvalves . . . . . . . . . . . . . . . . . . 2.3.3. Non-pneumatic membrane microvalves . . . . . . . . . 2.4. Passive mechanical microvalves for PCR chips . . . . . . . . . . 2.4.1. In-line polymerized gel microvalves . . . . . . . . . . . 2.4.2. Passive plug microvalves . . . . . . . . . . . . . . . . . 2.5. Passive non-mechanical microvalves -- hydrophobic microvalves 3. Fluid driving: micropumps for PCR microfluidic chips . . . . . . . . . 3.1. Mechanical micropumps for PCR chips . . . . . . . . . . . . . . 3.1.1. Piezoelectric micropumps . . . . . . . . . . . . . . . . 3.1.2. Pneumatic micropumps . . . . . . . . . . . . . . . . . 3.1.3. Thermopneumatic micropumps . . . . . . . . . . . . . 3.2. Non-mechanical micropumps for PCR chips . . . . . . . . . . . 3.2.1. Electrokinetic micropumps . . . . . . . . . . . . . . . . 3.2.2. MHD micropumps . . . . . . . . . . . . . . . . . . . . 3.2.3. Electrochemical micropumps . . . . . . . . . . . . . . . 3.2.4. Acoustic-wave micropumps . . . . . . . . . . . . . . . 3.2.5. Surface tension and capillary micropumps . . . . . . . . 3.2.6. Ferrofluidic magnetic micropumps . . . . . . . . . . . . 4. Fluid blending: micromixer for PCR microfluidic chips . . . . . . . . . 4.1. Passive micromixers for PCR chips . . . . . . . . . . . . . . . . 4.1.1. Y/T-type flow micromixers . . . . . . . . . . . . . . . 4.1.2. Recirculation flow micromixers . . . . . . . . . . . . . 4.1.3. Droplet micromixers . . . . . . . . . . . . . . . . . . . 4.2. Active micromixers for PCR chips . . . . . . . . . . . . . . . . 4.2.1. Electrokinetically-driven micromixers . . . . . . . . . . 4.2.2. Acoustically-driven micromixers . . . . . . . . . . . . . 4.2.3. MHD-driven micromixers . . . . . . . . . . . . . . . . 4.2.4. Micromixers with integrated microvalves/micropumps 5. Discussion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.1. Microvalves for microfluidic PCR chips . . . . . . . . . . . . . 5.2. Micropumps for microfluidic PCR chips . . . . . . . . . . . . . 5.3. Micromixers for microfluidic PCR chips . . . . . . . . . . . . . 6. Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Acknowledgements . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . for PCR chips . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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1. Introduction The polymerase chain reaction (PCR) (Saiki et al., 1985) is a key process in genetic analysis and has been playing an important role in modern biology and biochemistry research. However, the conventional 96well PCR devices usually have a thermal cycling rate of 1­2 °C/s, and therefore a complete PCR amplification needs 1­2 h or even longer. Moreover, by this approach, both the sample preparation and the post-PCR product detection need to be off-line performed, thus resulting in the longer analysis process and the higher risk of crosscontamination. Although the commercial capillarybased PCR systems can achieve the fast thermal cycling

rates of up to 10­20 °C/s (Walsh et al., 2005; Hataoka et al., 2005) and perform quantification analysis by analyzing the melting curves, it is difficult for them to perform the high-throughput nucleic acidic amplification. Furthermore, it is very inconvenient to manipulate the reaction capillary tube(s). The PCR mixture is first removed from each eppendorf and then pipetted into individual capillaries which are then capped. All capillaries are centrifuged to ensure that the PCR mixture fills the bottom of the capillary. The capillaries are then loaded into the capillary-based PCR device and undergo amplification. Nowadays, there have been many driving forces to exploit the potential benefits of micro-sized PCR

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apparatus (also called PCR microfluidic chip) relative to PCR systems of conventional size, including but not limited to reduced consumption of samples and reagents, shorter analysis times, greater sensitivity, portability, and disposability. The first chamber-based PCR chip devices were presented by Northrup and coworkers in 1993 (Northrup et al., 1993) and Wilding et al. in 1994 (Wilding et al., 1994). In 1998, Kopp and co-workers reported the first continuous-flow PCR chip (Kopp et al., 1998). These PCR chip models provided the basis for subsequent development, and up to now they have been replicated by numerous groups, sometimes with major amendments (Zhang et al., 2006). The PCR chips have undergone the transition from simple microfluidic components to highly integrated systems. Individual microfluidic components such as sample preparation, DNA/RNA amplification and/or product detection can be integrated on a single chip (Cady et al., 2005). Micropumps, microvalves and/or micromixers, especially dense pumping and valving arrays, can enable higher levels of integration, allowing a multitude of samples to be run in parallel so as to reduce analysis times and to reduce/avoid cross-contamination. In this review, we discuss the application of micropumps, microvalves and/or micromixers to the highly integrated PCR microfluidics since the first PCR chip was reported in 1993 (Northrup et al., 1993), and emphasize the pros and cons that exist in these functional components, indicating that the micropump, microvalve, and micromixer technologies have been playing a substantial role in witnessing the emergence of truly integrated PCR microfluidic chips that may have a markedly increased impact on the life and analysis sciences. 2. Fluid switching: microvalves for PCR microfuidic chips Microvalves are often one of the most important components for the realization of a fully integrated microfluidic system, such as lab-on-a-chip (LOC) or a micro total analysis system (TAS) (Manz et al., 1990; Auroux et al., 2002; Reyes et al., 2002; Vilkner et al., 2004; Dittrich et al., 2006). Not surprisingly, therefore, microvalves have been successfully applied to the field of the highly integrated on-chip PCR microfluidics. Nowadays, microvalves can be roughly classified into two groups: active microvalves, using mechanical and non-mechanical moving parts, as well as external systems, and passive microvalves, using mechanical and non-mechanical moving parts. In the following

sections, we will discuss the use of these micro-sized valve devices within the PCR chips in detail. 2.1. Active mechanical microvalves -- thermally actuated microvalves for PCR chips 2.1.1. Thermopneumatic microvalves Thermopneumatic microvalves are performed by volumetric thermal expansion coupled to membrane deflection. Typically, it has an actuator with a sealed chamber and movable diaphragm, where the liquid/gas or liquid-gas two-phase system is heated by the resistor incorporated in the chamber. The turn-on time depends on the heat mass and the available power of the heater, while the relaxation time is determined by the heat transfer to the external environment (Zdeblick and Angell, 1987; Zdeblick et al., 1994). By using this actuated microvalving technique, Baechi et al. presented a high-density, normally open thermopneumatic microvalve array for sample processing in PCR chips, as shown in Fig. 1 (Baechi et al., 2001). For this integrated chip, the in-line microvalve densities were significantly increased up to 300 valves cm- 2 by thermopneumatic actuation with a thin polydimethylsiloxane (PDMS) membrane and decreased thermal interaction by water cooling. The microchannel network with 5 m depth and 25 m width was designed for separation and combination of particles with a size between 0.1 and 5 m. The PDMS membrane, an area of 130 m × 30 m and a thickness of 2.9 m, was deflected up to 3.4 m with a heating power of 240 mW, and closing the microchannels needed 150 ms. The heat was released when the heater was switched off, which led to a fall time of 350 ms, which was longer than the rise time. Flow velocities in such a flow channel were 96 m s- 1 in an open mode and 8.3 m s- 1 in a closed mode. It is noted that the small size of this chip system integrated with aluminum thin-film heaters and temperature sensors allows thermal cycling with up to 40 °C s- 1. The desirable opening and closing of the array microvalves in the presented system can also be used for effective cell (or other bio-molecules) sorting, which has become an indispensable part in the studies of cell metabolism on a single-cell level (Fu et al., 2002). In addition, the use of PDMS membrane, an elastomer with low Young's modulus and high reversible strain, allows for large stroke and high sealing performance, as well as effective adherence between glass and silicon in the structure. 2.1.2. Shape memory alloy microvalves The shape memory effect is an attractive actuation principle for the development of microvalves, since it possesses simple and compact designs with high work

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stop valves without any external sensor for definite positioning of sample plugs (Fig. 2 (B)) (Münchow et al., 2005). Once the pressure-driven fluid plug passes the orifice, the driving pressure escapes through the hole and the fluid plug stops behind the ventilation hole at a defined position. After closing the vent by chipintegrated microvalves, the pressure on the fluid plug rises again and set the plug in motion. These valves are realized by a sealing piston operated by a SMA wire (or electromagnetic force). In this work, by using the presented microvalves and the ferrofluidic actuated micropumps (discussed below), the developed PCR chip was capable of processing all necessary steps such as sample preparation (including DNA extraction and preparation of the PCR mixture), merging and mixing, PCR as well as the final product detection. 2.2. Active non-mechanical microvalves -- phase change microvalves for PCR chips The phase change effect has been considered as an interesting new actuation technique for the development of microvalves. On the one hand, these microvalves appear in various configurations. Some use freeze water to create a plug, while others adopt paraffin, which is heated to open the microvalve (Felton, 2003). On the other hand, the solid-liquid phase transition can be easily adapted to a variety of lab-on-a-chip device for various applications including sealing, metering, and directed pumping. In addition, the microvalves based on this technique can be electronically addressable and effectively leak-proof, making them amenable for large-scale integration (Pal et al., 2004). In this section, we will introduce the applications of thermally-actuated, non-mechanical phase change microvalves to the fully integrated PCR chips. The phase change materials used in these microvalves include hydrogel, sol­gel, paraffin and ice. 2.2.1. Hydrogel microvalves Stimuli-responsive or smart hydrogen is able to change its volume reversibly and reproducibly by more than one order of magnitude even with very small alterations of certain environmental parameters. The volume change of smart hydrogels can respond to a variety of inputs including pH, glucose, temperature, electric field, light, carbohydrate and antigen (Oh and Ahn, 2006). More recently, Wang et al. have reported a normally-closed inline microvalve based on temperature sensitive hydrogels for sample flow control in the integrated PCR chips (Wang et al., 2005, 2006). The hydrogel microvalves were composed of a flow conduit, a cavity, and a hydrogel plug, as depicted schematically in Fig. 3 (A) (a). When the

Fig. 1. High-density thermopneumatic microvalve arrays for sample processing in PCR chips. (A) separation (A, B) and combination (B, C) of particle in a microchannel network; (B) Cross-section of a microvalve and a photodiode used for particle sensing in the channel network. Source: Baechi et al. (2001). [Fig. 1 (A) and (B) reprinted with permission from Baechi et al. (2001). Copyright (2001) Kluwer Academic Publishers.]

output, which are expected to allow a large miniaturization and operation at high pressure differences and flows. Kohl et al. developed a microvalve actuated by microfabricated shape memory alloy (SMA) thin films (NiTiPd), whose operation principle was shown in Fig. 2 (A). In the unheated martensitic phase condition, a pressure difference at the ports A and B causes a deflection of the membrane. Therefore, the microvalve is normally open (Fig. 2(A) (a)). By heating of the SMA microdevice above the phase transformation temperature, the undeflected shape is recovered and the membrane is pressed onto the valve seat as sketched in Fig. 2(A) (b). More recently, by using the SMA actuation method, along with the flow channel with hydrophobic vents, Münchow et al. reported a passive

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Fig. 2. (A) Operation principle of the SMA microvalves. Source: Kohl et al. (2000); (B) Principle of chip-integrated microvavles for sample exact positioning (A -- the pressure-driven fluid plug (blue) passes the hydrophobic ventilation hole (diameter = 800 m) and stops immediately after it; B -- the air can escape through the hole; C-After closing the hole by a piston the sample plug can move on). Source: Münchow et al. (2005). [Fig. 2 (A) reprinted with permission from Kohl et al. (2000). Copyright (2000) Elsevier Science Ltd.; (B) reprinted with permission from Münchow et al. (2005). Copyright (2005) Future Drugs Ltd.]

hydrogel was dry or when its temperature was increased above the critical temperature Tc (32 °C), the hydrogel remained in the de-swelling (unexpanded) state and the conduit was open to flow (top row in Fig. 3 (A) (a)). When the temperature decreased below the temperature Tc in the presence of aqueous solution, the hydrogel swelled and blocked the flow (bottom row in Fig. 3 (A) (a)). It was reported that the switching time of the hydrogel microvalve was 6 s for opening and 5 s for shutting off with the aid of electrically controlled thermoelectric (TE) units. In the open state, a flow rate of 84 L min- 1 was measured when the pressure drop across the valve was 6 kPa. In the absence of the hydrogel valve and under similar pressure head, the flow rate of the conduit was 1.9 mL min- 1. The closed valve did not exhibit any visible leakage up to a pressure of 200 kPa. In the first instance, Wang et al. used the microvalve of this kind to seal and to further pressurize the PCR chamber to minimize or eliminate the bubble formation that will adversely affect temperature uniformity and reduce the PCR amplification efficiency (Fig. 3 (A) (b)) (Wang et al., 2005). Subsequently, this temperatureactuated microvalve was incorporated to a pneumatically driven, disposable, highly integrated microfluidic cassette that comprised a PCR thermal cycler, an incubation chamber to label PCR products with up-converting phosphor (UPT) reporter particles, conduits, the hydrogel valves, and a lateral flow strip (Fig. 3 (B)) (Wang et al., 2006). The thermo-responsible hydrogel microvalves can be used for sample liquid distribution, metering, and the sealing of PCR chamber to suppress bubble formation. This kind of microvalves has many obvious advantages

such as simple fabrication and operation, perfect sealing, the ability to withstand relatively high pressure. Furthermore, in some cases, the microvalve can operate in a selfactuated, open-loop control mode, eliminating the need for a sensor to control the appropriate actuation time (Wang et al., 2005). 2.2.2. Sol­gel microvalves Pluronics sol­gel phase change microvalves for micro PCR were introduced by Liu et al. (2002c). The Pluronics sol­gel material is PCR-compatible, and 30% Pluronics polymer valves provide enough holding pressure of up to 138 kPa to ensure a successful PCR. The sol­gel polymer forms self-supporting cubic liquid crystalline gels at room temperature. By reducing the temperature locally to 5 °C, by a Peltier TE cooler unit, Pluronics microvalves were liquefied and opened. By using Pluronics sol­gel temperature transition microvalves, the integration of micro-PCR, biochannel hybridization, and subsequent hybridization wash functions in a single, low-cost, disposable monolithic device was accomplished, as shown in Fig. 4. The advantages of Pluronics phase change valves are their simplicity of implementation and operation. Although in solid gel form, Pluronics gel is not cross-linked and can be easily injected into the microfluidic valve structure to form a one-shot valve at room temperature. The Pluronics valves can be easily opened when valve temperature is lowered below the Pluronics gel transition temperature. However, the sol­gel based microvalves usually require liquid for actuation and are relatively slow.

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2.2.3. Paraffin microvalves Using a conventional solid material-paraffin for an active non-mechanical microvalve has been attractive due to the phase change nature of the material. The phase change material can be used either as a propellant for a membrane or as meltable plug to realize a microvalve. At present, two research groups reported thermally actuated paraffin microvalves as the meltable plug in microchannels to obstruct fluid flow in a PCR microfluidic chip: a reversible microvalve with external pneumatic air/vacuum systems (Pal et al., 2004, 2005) and an irreversible microvalve without external pneumatic air/ vacuum systems (Liu et al., 2004a,b). The plug changed phase from solid to liquid by thermal heating and moved in the microchannel by the external pneumatic air/vacuum system (Fig. 5 (A)) or by the pressure from the upstream liquid flow (Fig. 5 (C) (A­E)). The paraffin phase change plug is essentially leak-proof because of the phase change

Fig. 3. (A): (a) Illustration of the hydrogel valve's volume changes as the gel's temperature varies. The top (a) and bottom (b) rows correspond, respectively, to temperature above and below the critical temperature Tc. The left and right columns provide, respectively, top and side views. (b) A PCR chamber equipped with two hydrogel valves. Source: Wang et al. (2005); (B): (a) A photography of the microfluidic cassette comprising a PCR reactor, conduits, an incubation chamber and hydrogel microvalves. (b) A cartridge integrating the cassette (a) with the lateral flow strip. Source: Wang et al. (2006). [Fig. 3 (A) reprinted with permission from Wang et al. (2005). Copyright (2005) Springer Science + Business Media, Inc.; (B) reprinted with permission from Wang et al. (2006). Copyright (2006) Royal Society of Chemistry.]

Fig. 4. Monolithic integrated PC-based DNA assay device. Serpentine PCR channel (PCR), hybridization channel (HC), Pluronics valves (V1­V4), Pluronic traps (T), hydrophobic air-permeable membrane (M), PCR reagent loading holes (SL), sample driving syringe pump P1, waste-withdrawing syringe pump (P2), and wash syringe pump (P3). Source: Liu et al. (2002c). [Reprinted with permission from Liu et al. (2002c). Copyright (2002) American Chemical Society.]

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Fig. 5. (A) The reversible paraffin phase change microvalve with external air/vacuum systems. (i) Loading wax by actuating inlet port heater; (ii) Closing valve by actuating the inlet port and stem channel heaters with pressure at inlet port; (iii) Opening valve by actuating the stem channel and intersection heaters with vacuum at the inlet port. Source: Pal et al. (2004); (B) The integrated microfluidic device with reversible paraffin phase change microvalves. Source: Pal et al. (2005); (C) Schematic illustrations of a close-open paraffin microvalve design (A, B) and an open-close-open microvalve design (C­E). A photograph (F) of the PCR chamber surrounded by five paraffin based microvalves: valves 1­3 are open­close valves, and valves 4 and 5 are close­open valves. All valves are in "closed" position prior to initiating PCR. Source: Liu et al. (2004b). [Fig. 5 (A) reprinted with permission from Pal et al. (2004). Copyright (2004) American Chemical Society; (B) reprinted with permission from Pal et al. (2005). Copyright (2005) Royal Society of Chemistry; (C) reprinted with permission from Liu et al. (2004b). Copyright (2004) American Chemical Society.]

nature of the valve material: in the molten phase, the plug conforms to the channel walls and, once solidified, forms a solid seal. For the reversible microvalve, in the open state, the minimum pressure difference to induce motion of the water plug was less than 6.9 kPa; for the close state,

no leakage flows were observed over a period of 15 min up to 1725 kPa (Pal et al., 2004, 2005). For the irreversible microvalve the maximum hold-up pressure was about 275 kPa and no leakage was detected (Liu et al., 2004a,b). The time response of the paraffin microvalves is affected

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by the thermal actuation modes such as isothermal actuation, spatial or temporal gradient actuation (Pal et al., 2004). For the temporal operation mode of the reversible microvalve, the short actuation time (200 ms) was obtained, and no energy consumption was required to maintain either the open or the closed state, making the valves extremely energy efficient. However, in the irreversible paraffin microvalve, the time required to open and close valves was on the order of tens of seconds (up to 15 or 20 s) (Liu et al., 2004a). Although the time response of these valves is relatively slow as compared to most conventional microvalves (ms), the paraffin-based microvalves are practical and useful in some microfluidic applications where rapid valving time is not critical, such as micro PCR chips (see Fig. 5 (B) and (C) (F)). It is noted that the time response of the paraffin valves is also dependent on the heating rate of the heater, as well as on the physical nature of the channel substrate material (i.e. thermal mass and thermal conductivity). Another approach to improve the time response is to use a phasetransition material with a lower melting temperature. 2.2.4. Ice microvalves Microvalves can also be formed by ice plugs that noninvasively close and open the microchannel and, thus, prevent or enable the sample liquid flow through the microchannel. Gui and Liu designed an ice microvalve by using a TE cooling and heating unit that was in contact with a microtube. The microvalve was opened and shut off by thermally-actuated phase change of the aqueous solution, i.e., melting and freezing, within several seconds (Gui and Liu, 2004). In 2001, in fact, He et al. demonstrated the feasibility of freeze-thaw microvalving in a capillary-based fully integrated and automated PCR system (He et al., 2001). More recently, in a similar approach, Chen et al. applied this technique to a polycarbonate (PC) microchip comprising a more complicated microfluidic channel structure and demonstrated PCR inside an ice microvalve-controlled stationary chamber (Chen et al., 2005b). It is noted that the ice microvalving technology can also be conveniently utilized for sample metering and distribution that takes advantage of the self-actuated mode of valve operation, and for filling and withdrawing samples from a closed loop (Chen et al., 2005b). The response time of ice microvalve is depended by many factors and can be significantly shortened in certain situations. In the work by Chen et al., the response time of ice valve ranged from 0.5 s (under pre-cooled conditions) to about 10 s (under high flow conditions and without precooling). The ice microvalves are the closest in concept to the paraffin microvalves but have the advantage of being

normally opened at room temperature, which is especially advantageous in systems in which liquid plugs are pneumatically driven. In addition, although in direct contact with the sample fluid, the ice valves using the aqueous solution as the valving medium won't raise concerns of contamination and biocompatibility. 2.3. Active external microvalves for PCR chips In this section, previous work on active microvalves using external systems in the fully integrated PCR chips will be discussed. They include modular built-in valves, thin membrane or in-line microvalves actuated by external pneumatic air pressure or vacuum. The nonpneumatically-actuated membrane microvalves will also be described. Using external systems is one of the most practical approaches in devising microvalves. This actuation can afford advantage with no leakage at high input pressures, although miniaturization may be difficult due to the requirement of additional external powering and pumping systems. 2.3.1. Modular built-in microvalves Oh et al. developed a world-to-chip microfluidic interfacing system with modular built-in microvavles with no dead volume, no leakage flow, and biochemical compatibility, as shown in Fig. 6 (Oh et al., 2005a,b). This world-to-chip microfluidic valves performed excellently in both loading the samples and sealing the reagents, as evidenced in successfully performing multichamber real-time PCR amplifications with no contamination or leakage failures. Valving of the inlet, outlet or vent ports is a critical function to successfully carry out PCR. The microvalves withheld an internal pressure of about 47 kPa generated all through thermal cycling and were reopened easily after PCR. 2.3.2. Pneumatic microvalves 2.3.2.1. Membrane microvalves. Membrane-based pneumatic microvalves were made of a laminate with a mylar layer sandwiched between two silicone layers (Anderson et al., 2000), thin latex sheets (Lagally et al., 2000, 2001a,b), thin PDMS layers (Lagally et al., 2003, 2004; Toriello et al., 2006; Liu et al., 2006; Easley et al., 2006; Liao et al., 2005; Lien et al., 2007; Huang et al., 2006; Hou et al., 2007) or thin 3 M tape sheets (Yuen et al., 2000). The appropriate application of this kind of membrane microvalves to a highly integrated microfluidic biochip was first presented by Anderson et al. in 2000 (Anderson et al., 2000). This presented microdevice was capable of extracting and concentrating nucleic acids from

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Fig. 6. Schematic views of the world-to-chip microfluidic interfacing approach with modular built-in valves for both sample loading and sample sealing. (A) Sample loading mode; (B) Sample sealing mode. Source: Oh et al. (2005a). [Fig. 6 (A) and (B) reprinted with permission from Oh et al. (2005a). Copyright (2005) Royal Society of Chemistry.]

milliliter aqueous samples and performing microliter chemical amplification (including reverse transcription (RT), PCR and nested PCR), serial enzymatic reactions, metering, mixing and nucleic acid hybridization. Diaphragm valves were closed and opened using 345 kPa and a vacuum, respectively, and they were used in conjunction with fluid barriers and gas pressure to position and manipulate reagents. Each barrier was a porous hydrophobic membrane that allows the passage of gas but not of aqueous liquid, and a reagent plug guided into the chamber will stop moving when it reaches the fluid barrier. This approach was used to position and to meter reagents without using sensors or feedback. In addition to guiding each reagent plug to the appropriate chamber, the diaphragm valves physically prevent reagent migration of evaporation during high temperature amplification reactions (Anderson et al., 2000). Using a design modeled after that presented by Anderson et al. (Anderson et al., 2000), Mathies's group have made attempts to develop practical pneumatic microvalves with latex membranes (Lagally et al., 2000, 2001a,b) for fully integrated PCR-capillary electrophoresis (CE) microfluidic devices as depicted in Fig. 7. Two aluminum manifolds, one each for the hydrophobic vents and latex membrane valves, were placed onto the respective ports and clamped in place using vacuum. The manifolds were connected to external solenoid valves for pressure and vacuum actuation. Sample was loaded from a right inlet port by opening the valve using a vacuum (4 kPa) and forcing the sample under the

approximately 150 m thick latex membrane using pressure (69­83 kPa). Vacuum was simultaneously applied at the hydrophobic vent to evacuate the air from the chamber. The sample stopped at the vent, and the valve was pressure-sealed to enclose the sample. Although this approach has notable advantages such as sensorless sample positioning to 200 nL PCR chambers and bubble evacuation, the latex microvalving technology has large 50 nL dead volumes per element and requires time-consuming manual construction. (Lagally et al., 2003, 2004). In order to circumvent this issue, this research group brought forward an improved method for constructing large arrays of polydimethylsiloxane (PDMS) membrane microvalves on glass microdevices (Grover et al., 2003), as shown in Fig. 8. The dead volumes for these PDMS valves can be as small as 8 nL and they can be actuated with small pressures and vacuums. This PDMS valve technology has been successfully combined with the PCR-CE chip to form a multi-layer glass microdevice capable of amplifying and electrophoretically analyzing the PCR product (Lagally et al., 2003, 2004; Toriello et al., 2006; Liu et al., 2006, 2007). In addition to the dead volumes as small as 8 nL and small pressures and vacuums for actuation, three PDMS membrane microvalves in series formed versatile membrane diaphragm pumps. In addition, these microvalves are normally closed and therefore minimize PDMS solution contact to avoid the well-known chemical absorption and fluorescence background problems of PDMS. Pneumatically-actuated PDMS membrane microvalves have also been reported by other groups (Easley et al., 2006; Liao et al., 2005; Huang et al., 2006; Lien et al., 2007; Hou et al., 2007). Landers' group has recently utilized the normally-closed PDMS microvalves developed by Grover et al. (Grover et al., 2003) for on-chip pressure injection for integration of infrared-mediated non-contact DNA amplification with electrophoretic separation of the products (Easley et al., 2006). A commercially available PDMS membrane with a thickness of 254 m was used as the deflectable valve layer. An oil-less diaphragm vacuum pump/compressor was used to control the pneumatic valve lines by application of pressure (15 kPa) to keep valves closed or vacuum (60 kPa) to open them. Actuation of these valves was accomplished using solenoid valves and corresponding manifold. The main advantages pertinent to the presented microvalve configuration include as follows: (i) the valves were pneumatically addressable and effectively leak-proof, making them amenable for present large-scale integration; (ii) the use of the normallyclosed PDMS valves for this integrated process design allowed individual regions of the chip to be completely

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Fig. 7. The PCR-CE microfluidic device (A) with pneumatically actuated membrane microvalves and hydrophobic vents (B). Source: Lagally et al. (2001b). [Fig. 7 (A) and (B) reprinted with permission from Lagally et al. (2001b). Copyright (2001) American Chemical Society.]

Fig. 8. Cross-section and top views of three-layer (A) and four-layer (B) monolithic pneumatic PDMS membrane microvalves. Source: Grover, et at. (2003). [Fig. 8 (A) and (B) reprinted with permission from Grover et al. (2003). Copyright (2003) Elsevier Science Ltd.]

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isolated either before or after attaching the valve layers, making them convenient to experimentally operate and run; (iii) the valve actuation channels were arranged in a trifurcated pattern at the point of connection to the valve seats, allowing for more even actuation of the valve seats when compared to simple one-channel actuators; (iv) finally the valves were arranged to allow diaphragm pumping from the PCR chamber, the marker reservoir, or from both into the CE separation region. Lee's group also focused on the development of integrated PCR chips with PDMS membrane microvalve(s) and PDMS membrane micropumps resulting from the

peristaltic effect by the individual membrane microvalve (Liao et al., 2005; Huang et al., 2006; Lien et al., 2007). Liao et al. developed a miniaturized RT-PCR system for diagnosis of RNA-based viruses, in which the reagent and sample flow was controlled automatically by means of two micropumps and one microvalve, as shown in Fig. 9 (A) (Liao et al., 2005). The fluid flow could be successfully interrupted when the pneumatic microvalve was closed, and the pressure of 69 kPa could cause the microvalve to close the flow of a fluid moving at 30 L min- 1. Using the similar microvalving and micropumping approach, subsequently, this research group reported an integrated

Fig. 9. (A) Schematic diagram of two-step micro RT-PCR chip integrated with micropumps and microvalves for controlling biosample transportation. Source: Liao et al. (2005); (B) Schematic illustration of a microfluidic chip capable of DNA/RNA amplification, electrophoretic injection, separation, and on-line detection of product detection. Source: Huang et al. (2006); (C) Schematic illustration of the integrated RT-PCR chip. Several components including a microtemperature module, a bead collection module, and a microfluidic control module are integrated onto a single chip. Source: Lien et al. (2007). [Fig. 9 (A) reprinted with permission from Liao et al. (2005). Copyright (2005) Oxford University Press; (B) reprinted with permission from Huang et al. (2006). Copyright (2006) WILEY-VCH Verlag GmbH & Co. KGaA; (C) reprinted with permission from Lien et al. (2007). Copyright (2006) Elsevier Science Ltd.]

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microfluidic chip for DNA/RNA amplification, electrophoresis separation and on-line optical detection by a buried optical fiber, as depicted in Fig. 9 (B) (Huang et al., 2006). Three micropumps consisting of three PDMS membranes were designed. When one micropump was activated, the other two served as microvalves simultaneously. Therefore, the PDMS membranes could provide a valving function to ensure a proper isolation of the PCR reagents and the CE buffers. Most recently, this group developed an integrated RT-PCR microfluidic system for virus detection (Lien et al., 2007). Comparable to a largescale apparatus, this integrated microfluidic chip could perform mixing, incubation, purification, transportation, and nucleic acid amplification of virus automatically with the aid of integrated micropumps and microvalves, as shown in Fig. 9 (C). In these PDMS valving systems, the pressure of up to 173 kPa was used for the operation of the microvalves to shut off the fluid flow, while at this pressure the maximum peristaltic pumping rate was measured to be 100 L min- 1 at the driving frequency of 11 Hz. Here, it needs to be noted that a semi-disposable CNC (computer numerical control) machined plexiglass microvalve system with flexible 3M tape membranes was presented by Yuen et al. (2000). The microvalve, designed for human cell isolation and DNA amplification systems, could process microliter and sub-microliter samples of human whole blood with minimal dead volume of less than 0.16 L and maximum sealing pressure up to 2069 kPa, and without attendant problems such as loss of sample on the microvalve system and loss by evaporation. The advantage of the CNC machined plexiglass microvalve is that it is easy to manufacture and it provides an alternative for the conventional silicon chip-based microvalve systems. 2.3.2.2. In-line microvalves. Quake's group reported a series of microfluidic chip systems by using the multilayer soft lithography technique (Unger et al., 2000; Quake and Scherer, 2000; Thorsen et al., 2002; Fu et al., 2002; Chou et al., 2001; Liu et al., 2002a, 2003b). A crossed-channel building block in the microfluidics systems was a pneumatically actuated in-line microvalve as shown in Fig. 10 (A) and (B). It consisted of the upper channel ("control channel"), the lower channel ("flow channel"), and the membrane of polymer between the channels. Two layers formed by the PDMS rapid prototyping technique were bonded together in a crossed-channel architecture (Unger et al., 2000; Quake and Scherer, 2000). The bonded structure was sealed onto the top of a glass or elastomer substate. When pressurized gas was applied to the upper pneumatic control channels, the rubber membrane deflected at the intersection of the in-line

microchannels at the lower layer. Typical channels were 100 m wide and 10 m high, making the valve an active area of 100 m × 100 m. The microvalve was closed with a pneumatic pressure of 100 kPa and back to its rest opening position by its own restoring spring force (40 kPa), with the dead volume of 100 pL and the time response on the order of 1 ms (Unger et al., 2000). Quake's group has utilized the concept of these pneumatically actuated in-line microvalves for the PCR microfluidic systems (Chou et al., 2001; Liu et al., 2002a, 2003a) due to ease manipulation of nanoliter sample volumes in the in-line microchannels. For example, Liu et al. reported a nanoliter rotary microfluidic device for PCR amplification (Chou et al., 2001; Liu et al., 2002a). Such a device consisted of two layers: fluid channels on the bottom and pneumatic actuation channels on the top. The bottom fluid layer had two sample inputs, a mixing T-junction, a central circulation loop, and an output channel. The top layer had several stand-alone actuation channels, which could be pressured or vented to atmosphere. Any intersection of a top air channel with a bottom fluid channel formed a microvalve. The on-chip microvalves were actuated with an external pneumatic controller that allowed selective actuation of valves to seal off the loop as well as peristaltic pumping at variable rates within the loop, and typical actuation pressures were 69 kPa (Liu et al., 2002a). Large-scale integration (LSI) in microfluidic systems analogous to that in electronic integrated circuits was realized by building up each in-line microvalve (Thorsen et al., 2002). As an example of the microfluidic LSI, a 20 × 20 microfluidic channel matrix requiring merely 41 pipetting steps for 400 distinct PCR reactions was demonstrated (Liu et al., 2003a), as depicted in Fig. 10 (C). A single 2 L aliquot of DNA polymerase was distributed over all 400 independent reactions (3 nL). In total, 2860 in-line microvalves displayed horizontally or vertically were controlled by only two independent pneumatic pressure supplies. Furthermore, the large valves or the small valves were selectively actuated because they had different thresholds of hydraulic pressure necessary for actuation. The microfluidic LSI matrix chip provided a universal and flexible method for biological and chemical assays with low consumption of precious reagents in a highly automated fashion. 2.3.3. Non-pneumatic membrane microvalves Almost all of the forementioned PDMS-based microvalves are actuated by external pneumatic pressure manifolds. However, the external components required for such an approach (for example, high pressure gas systems

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Fig. 10. (A) Process flow for multilayer soft lithography; (B) Schematic of pneumatically-actuated in-line microvalve closing for square and rounded channels. The dotted lines indicated the contour of the top of the channel for rectangular (left) and rounded (right) channels as pressure was increased. Source: Unger et al., 2000. (C) Schematic diagram of the N = 20 matrix chip to perform 400 independent PCR reactions, with in total 2860 in-line microvalves that was controlled by only two independent pneumatic pressure supplies. Source: Liu et al. (2003a). [Fig. 10 (A) and (B) reprinted with permission from Unger et al. (2000). Copyright (2000) American Association for the Advancement of Science; (C) reprinted with permission from Liu et al. (2003a). Copyright (2003) American Chemical Society.]

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and their valves) result in a relatively complex system. The following examples are used to illustrate that the nonpneumatically actuated PDMS diaphragm microvalves can be desirably used for the PCR microfluidic chips. Backhouse's group developed an adaptable microvalving system based on the use of three servomotor-controlled valve fingers that actuated PDMS microchannels in the PCR microfluidic chip, as shown in Fig. 11 (Pilarski et al., 2005). The shut-off of fluid flow in the microchannel was accomplished by lowering one of the outer valve fingers onto the PDMS and applying pressure until the channel was fully compressed, and complete compression could be inferred automatically from power consumption of the servomotor or through visual inspection. The servo current draw between 170 mA and 180 mA corresponded to a complete channel closure, with little fluid loss from the PCR chamber and no visible damage to the PDMS surface. These non-pneumatic fingerlike microvalves can allow a single valving system to be used for multiple runs on multiple chips, reducing fabrication cost of a single chip and allowing chip prototypes to evolve independently of the valving system. In addition, this valve approach allows for a very compact valving system and facilitates higher valving densities and resolution by using valve fingers with a smaller radius of curvature. Also, noteworthy is that within the PCR chip the about 1 mm thick PDMS roof structure of the PCR chamber can also serve as the simple peristaltic diaphragm pumping actuation, which simplifies the chip design and reduces the fluid dead volume in the chip. Recently, Backhouse's group has applied this servomotor-driven microfluidic valving and pumping technology to other PCR chip systems, with integrated fluid control and vapour barrier (Prakash et al., 2005) or

with integration of small volume PCR amplification and CE separation detection to assess risk of BK virusassociated nephropathy in renal transplant recipients (Kaigala et al., 2006). 2.4. Passive mechanical microvalves for PCR chips In this section, we will briefly review the applications of the passive mechanical microvalves to the integrated PCR chips. Compared with most of passive mechanical microvalves such as flaps, membranes, spherical balls or mobile structures, the passive mechanical microvalves discussed in this section have been rarely reported in many fully integrated microfluidic devices, although they have been successfully integrated onto the PCR chips. 2.4.1. In-line polymerized gel microvalves Fan's group utilized the in situ gel photopolymerization method to create local gel plugs that function as passive microvalves effectively preventing bulk flow of liquid due to the thermally induced pressure differences during thermal cycling. After PCR, the negatively charged products were electrokinetically driven through the gel plug, followed by on-chip CE separation detection (Fan et al., 2003; Koh et al., 2003). The used photoinitiator was 1-hydroxylcyclohexylphenylketone (HCPK), and the solution for fabricating microfluidic gel valves consisted of Tris­HCl, acrylamide/bic and HCPK. These reported gel microvalves could withstand hydrostatic pressures up to 690 kPa without moving and with minimal leakage. However, in these gel microvalves, the degree of gel cross-linking has a large effect on the leakage-proof performance of the microvalve: the degree of leakage in 4% cross-linked gel is about twice that in 10% cross-linked gel, and thus the preparation of gel plug needs to be precisely and carefully performed. In addition, there will be a bias in the amount of DNA sample electrokinetically injected through the gel microvalve if the injection time is too short. That is to say, if the bias in the amount of injected samples needs to be reduced or eliminated, the injection time of samples through the gel valve must be long enough. This process is disadvantageous to the fast DNA analysis. 2.4.2. Passive plug microvalves In order to implement a practical integrated PCR-CE microfluidic chip, Prakash and Kaler have most recently presented a passive plug (PP) microvalve fabricated using PDMS elastomer that was incorporated in the integrated chip, as shown in Fig. 12 (A) (Prakash and Kaler, 2007). The PP microvalves with the snug and airtight stub and stem components utilized an airtight

Fig. 11. Schematic representation of the valving finger assembly actuated by three servomotors. Source: Pilarski et al. (2005). [Reprinted with permission from Pilarski et al. (2005). Copyright (2005) Elsevier Science Ltd.]

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plugging mechanism to enable fluid/vapor retention, and sample transport from each PCR chamber to the corresponding CE sample well. In addition to the snug fit, application of thrust (e.g. by weight) above the stub of the PP microvalves will also help a tight seal inside the holes. It is noted that the PP microvalves and discontinuous channel layout with a `flow control elevation' inside the access port could be used to control fluid flow. By using these PP microvalves and their ability to pump sample fluids, the nine-array high-throughput PCR-CE chip has been successfully developed and tested, as depicted in Fig. 12 (B). To the best of our knowledge, the throughput degree of this PCR-CE chip is the highest up to now, compared with other reported PCR-CE chips (Lagally et al., 2000, 2001a,b, 2004; Kaigala et al., 2006; Liu et al., 2006, 2007; Huang et al., 2006; Easley et al., 2006; Koh et al., 2003; Toriello et al., 2006). Although the operation of these PP microvalves is simple, efficient and they are mechanically robust, the direct contact of the PP valve components with PCR sample fluid will result in the possible cross-contamination. In addition, the operation of these PP microvalves is still manual, and thus the degree of automation is low. 2.5. Passive non-mechanical microvalves -- hydrophobic microvalves for PCR chips In PCR microfluidic chips, an important approach to controlling fluid flow is the passive non-mechanical hydrophobic microvalves utilizing the surface properties in the microchannels (Gong et al., 2006; Münchow et al., 2005; Liu et al., 2002c; Lagally et al., 2000, 2001a,b; Anderson et al., 2000; Burns et al., 1998). A hydrophobic patch involving hydrophobic regions in otherwise hydrophilic microchannels can be desirably used as a microvalve. For example, Gong et al. used a hydrophobic chemical (EGC-1700) to selectively treat the air venting channel surface inside the glass chip so as to form a hydrophobic patch valve confining the PCR mixture within the microchamber (Gong et al., 2006). Certain chips could be equipped with access ports that were connected to the main channel for definite positioning of fluid plugs. By coating these ports with hydrophobic Teflon substances the structures could be used as passive stop valves without any external sensor, as shown in Fig. 2 (B) (Münchow et al., 2005). Liu et al. utilized an air-permeable hydrophobic membrane to seal an access hole at the end of the hybridization channel in the PCR-DNA hybridization chip, as demonstrated in Fig. 4 (Liu et al., 2002c). The similar works have also been reported by other groups (Lagally et al., 2000, 2001a,b; Anderson et al., 2000; Burns et al., 1998). Although the hydrophobic micro-

valves have been desirably applied to the integrated PCR chips, to stably seal the PCR mixture inside the microchamber with these hydrophobic microvalves is a challenging matter. On the one hand, the stability of the hydrophobic film coated inside the microchannels is problematic; on the other hand, the hydrophobic microvalves usually withstand the low pressure. Also, the surfactants such as polyethylene glycol (PEG), polyvinylpyrrolidone (PVP) and Tween 20, which are often included into the PCR mixture to stabilize the PCR amplification, maybe will destroy the hydrophobic nature of these microvalves. 3. Fluid driving: micropumps for PCR microfluidic chips For many microfluidic systems, a self-contained, active micropump, whose packaging size is comparable to the fluid volume to be pumped, is necessary or highly desirable. Most probably inspired by the beautiful application future of LOC or TAS for chemical and biological analysis (Manz et al., 1990; Auroux et al., 2002; Reyes et al., 2002; Vilkner et al., 2004; Dittrich et al., 2006), a diverse series of micropumps have been reported for the development of highly integrated, (quasi-)automatically-operated PCR chips. In the following sections, we will review the applications of some microfluid pumping strategies to the PCR chips. Most micropumps found today can be roughly categorized into two groups: mechanical micropumps with moving parts and non-mechanical micropumps without moving parts. Two movement mechanisms have been employed in mechanical micropumps: reciprocating and peristaltic movements. According to various actuation principles, the mechanical micropumps can be subdivided into several categories: piezoelectric, pneumatic, thermopneumatic, electrostatic, electromagnetic and bimetallic SMA micropumps, while the non-mechanical micropumps mainly include electrokinetic, magnetohydrodynamic (MHD), electrochemical, acoustic-wave and surface tension and capillary micropumps. 3.1. Mechanical micropumps for PCR chips 3.1.1. Piezoelectric micropumps The use of piezoelectrics to drive micropumps can be traced back to the early 1970's (Demer, 1974), but the publications of Van Lintel and Smits on diaphragm and peristaltic micropumps with piezoelectric actuation (Van Lintel et al., 1988; Smits, 1990) mark the beginning of extensive micropump research in the microelectromechanics systems (MEMS) world. Nowadays, many piezo-

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Fig. 12. (A) Overview and functionality of the passive plug (PP) microvalves. (B) Picture of the nine-array PCR-CE glass microfluidic chip integrated with the PP microvalves. Source: Prakash and Kaler (2007). [Fig. 12 (A) and (B) reprinted with permission from Prakash and Kaler (2007). Copyright (2006) Springer-Verlag.]

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driven micropumps with varying valve types, geometries and fabrication technologies have been developed and applied in many different fields, but piezoelectric actuators are employed in a similar manner (Laser and Santiago, 2004; Woias, 2005). In 2003, Bu et al. from University of Southampton (UK) successfully used this type of microfluidic pumping technique to realize the micro oscillatoryflow PCR chip (Bu et al., 2003). A bidirectional peristaltic pump, which was driven by three piezoelectric discs located in a recess etched in the glass, was connected to the third reaction chamber by a channel. Numerical simulation was used to obtain the optimum pumping performances. By optimizing the pump membrane parameters, the presented piezoelectrically-actuated micropump could provide a stroke volume of 314 nL. A maximum pump rate of 3.14 L s- 1 with an operating frequency of 10 Hz was obtained, and a maximum pressure differential that the micropump could withstand was 50­60 kPa. The flow rate of 3.14 L s- 1 allowed a PCR mixture of 1 L to be pumped between three adjacent chambers in 0.6 s through a connection channel of 1 L (Bu et al., 2003). Although the piezoelectric actuation is a very attractive concept since it provides a comparatively high stroke volume, a high actuation force, and a fast mechanical response, it hasn't been widely used for the development of fully integrated PCR chips up to now. However, we believe that it will soon be routinely implemented for fluid manipulation within the PCR chips. It is noted that commercial piezoelectric transducer (PZT) material is readily available for a hybrid integration. 3.1.2. Pneumatic micropumps As discussed before, the pneumatically-actuated microvalves have been widely in the integrated PCR chips, compared with other types of microvalves. When a series of on/off actuation sequences are applied to three active pneumatic microvalves, a pneumatic peristaltic micropump will be formed, as shown in Fig. 13 (A). If these pneumatic microvalves are placed in series, a pneumatically-actuated diaphragm micropump will be formed, as depicted in Fig. 13 (B). As with active external microvalves mentioned above, the PDMS elastomer materials of low modulus are widely used in pneumatically-driven peristaltic (Liao et al., 2005; Lien et al., 2007; Huang et al., 2006; Chou et al., 2001; Liu et al., 2002a, 2003a) or diaphragm (Easley et al., 2006) micropumps. In the pneumatic peristaltic micropumps, three separate microvalves are usually used to peristaltically pump the working fluid (Liao et al., 2005; Lien et al., 2007; Huang et al., 2006). However, one S-shaped channel could accomplish the same peristaltic pumping as three separate channels, and provided more secure closing of the inlet and

Fig. 13. (A) A 3D scale diagram of a pneumatic peristaltic micropump. Source: adapted from Unger et al. (2000); (B) An oblique view of the three-layer pneumatic diaphragm micropump. Source: adapted from Grover et al. (2003). [Fig. 13 (A) reprinted with permission from Unger et al. (2000). Copyright (2000) American Association for the Advancement of Science; (B) reprinted with permission from Grover et al. (2003). Copyright (2003) Elsevier Science Ltd.]

outlet than a single channel (Liu et al., 2002a). In a similar approach, instead of using a traditional three-membrane pneumatic pump requiring three electromagnetic valves (EMVs), Lien et al. recently reported a new pneumatic peristaltic micropump requiring only one EMV, as shown in Fig. 9(C). The compressed air filled up the cavities in subsequence such that the sample solutions could be pushed forward (Lien et al., 2007). This new peristaltic micropump obtained the maximum fluid pumping rate up to 100 L min- 1 at the driving frequency of 11 Hz and at a pressure of 173 kPa, which was much higher than those obtained at the same pumping conditions in the traditional three-membrane pneumatic mciropump (Liao et al., 2005; Huang et al., 2006). Whether the pneumatic peristaltic or diaphragm micropumps, they require an external pneumatic supply and one or more high-speed valve connections, and they therefore are not strictly comparable to micropumps with fully integrated actuators such as the forementioned piezoelectric micropumps. In order to overcome this, Backhouse's group has more recently reported a kind of electric­servomotor-driven peristaltic micropump for sample fluid pumping in the integrated PCR chips, in which the pumping operation was obtained through the sequential actuations of three finger-shaped rods, as shown in Fig. 14 (Pilarski et al., 2005; Prakash et al., 2005; Kaigala et al., 2006). This concept of micropump is simple, effective, and compact relative to most macroscopic

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pumps and pneumatically-actuated micropumps. However, these micropumps are still larger than the micropumps integrated with miniature actuators. It needs to be noted, in addition, that although the peristaltic or diaphragm micropumps discussed here are used in unidirectional mode, they can also allow bidirectional sample fluid transport by simply changing the microvalve actuation scheme. 3.1.3. Thermopneumatic micropumps The thermopneumatically driven micropump was first demonstrated by Van de Pol et al. in 1990 (Van de Pol et al., 1990), where an air-filled chamber with an internal heater resister was incorporated on top of the pump diaphragm. Although this type of micropump can be made very compact, fabricated using standard micromachining processes, and operated at low-voltage operating conditions, they usually have a low pump rate, and a relatively long temporal response limited by the rate of heat transfer, especially during the cooling process. Moreover, they still belong to the membrane-actuated micropumps and they therefore suffer from complicated design, complicated fabrication, or high cost. In order to circumvent this shortcoming, Liu et al. developed a simple thermopneumatic air micropump without membrane actuation for the self-contained, fully integrated PCR chip (Liu et al., 2004b), as shown in Fig 15. This micropump made use of the air expansion in an air chamber, which was attached to a resistive heater, when heated. The air expansion was a

nearly linear function of temperature. The resulting air expansion pushed the solution from the storage chamber into the downstream channels and chambers. The presented air micropump with air pockets of 50 L internal volume could efficiently move up to 60 L volume of fluids with a heater power consumption of less than 0.5 W. In a different approach, Burns's group first trapped the gas in pockets behind the sample and then heated them so as to drive sample and reagents (Burns et al., 1998; Mastrangelo et al., 1999). This type of thermopneumatically driven micropump is sometimes referred to as the `bubble' pump, in which pumping is driven by phase change of the primary working fluid, rather than of a secondary working fluid in a separate chamber. 3.2. Non-mechanical micropumps for PCR chips It is possible to move liquids and constituents in liquids without using any moving parts. These techniques use a variety of interaction phenomenon between an electromagnetic field and the working fluid to generate pressure and flow. 3.2.1. Electrokinetic micropumps The actuation principle of electrokinetic micropump is based on the movement of molecules in an electric field due to their charges, which is widely used to move liquids and particles in microchannels. There are two components to electrokinetic flow: electrophoresis and electroosmosis. The former results from the accelerating force due to the charge of a molecule in an electric field balanced by the frictional force. The latter leverages the surface charge that spontaneously develops when a liquid comes in contact with a solid. A prerequisite for the electroosmosis is the presence of immobilized surface charges at the microchannel wall in contact with an electrolyte solution. This surface charge leads to the formation of an electric double layer (EDL) by attracting oppositely charge ions from the buffer and therefore leads to concentration and charge density gradients in the immediate vicinity of the wall. Electrokinetic micropumping is the basis of micro capillary electrophoresis separation and a comprehensive treatment of this subject would by far exceed the scope of this review. Here, we only introduce the desirable applications of electrokinetic micropumps for sample fluid driving in the PCR chips (Lee et al., 2005; Chen et al., 2005a; Hu et al., 2006; Gui and Ren, 2006). Lee et al. used an electroosmotic micropump for sample transport in the PCR microchannel, where Au microelectrodes were deposited on the glass substrate as the driving electrodes for electroosmosis flow (EOF) pumping, and tested the effect of EOF driving electric field on the on-chip PCR products (Lee et al.,

Fig. 14. Sequence of the valving finger actuation during a pumping circle. The fingers were driven by three electric servomotors. Source: Pilarski et al. (2005). [Reprinted with permission from Pilarski et al. (2005). Copyright (2005) Elsevier Science Ltd.]

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Fig. 15. Schematic of the fully integrated PCR chip with a thermopneumatic air micropump (#4) and three electrochemical pumps (#1­3). Source: Liu et al. (2004b). [Reprinted with permission from Liu et al. (2004b). Copyright (2004) American Chemical Society.]

2005). The electrokinetic micropumps have been well developed for the continuous-flow PCR chips with varying cycle number. For example, Chen et al. have recently developed an electrokinetically synchronized PCR microchip fabricated in PC, as shown in Fig. 16 (Chen et al., 2005a). The synchronized format allows fabricating a shorter length microchannel for the PCR compared to nonsynchronized continuous flow formats, permitting the use of smaller voltages when the flow is driven electrically and also allows flexibility in selecting the cycle number with no need to change the chip architecture. Hu et al. used the electrokinetic pumping technique to realize the twotemperature oscillating-flow PCR thermal cycling chip, where the PCR mixture was electrokinetically pumped through the microchannel forwards and backwards, with a higher current for higher temperature and a lower current for lower temperature (Hu et al., 2006). Most recently, Gui and Ren have used the numeric simulation approach to demonstrate the possibility of employing electrokinetic flow for a unidirectional PCR chip design with serpentine microchannel with fixed cycle number (Gui and Ren, 2006). They investigated the effects of multiple controlling parameters, such as the layout of microchannels, channel geometry, applied electrical field strength, or chip substrate

material, on eletroosmosis-based continuous-flow PCR processes. Although the electrokinetic micropumping technology has been desirably used in the PCR chips, it has its pros and cons. Electrokinetic flow has the flat and pluglike velocity profile reducing sample dispersion. Moreover, if using electrokinetic flow, no moving element is required in the system, and electric control makes the system easy to operate and integrate. Nevertheless, electrokinetic flow has important drawbacks for PCR bioassays. On the one hand, the velocity of electrokinetic flow is relatively slow, compared with that of pressuredriven flow, leading to the PCR processing times longer. On the other hand, electrophoretic demixing -- the separation of components in a heterogeneous mixture due to different electrophoretic mobilities is unfavorable in PCR assays. A continuous-flow PCR assay often requires a uniform flow for all species, especially in PCR amplifications in which some additives are included for dynamic passivation in order to reduce the adsorption of PCR reagents onto the chip inner surface as a result of high surface-to-volume ratio. Also, sample purity is another problem special attention should paid to, because adsorption of proteins and other impurities will perturb

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with a MHD micropump (West et al., 2002). The annual microchannel for PCR reactions capable of MHD actuation was made of silicon or SU-8, and the electric field used was an AC electric field. With a miniature electromagnet producing a magnetic flux density of 7.4­ 13.5 mT and a peak channel current between 55 and 156 mA, this micropump could produce an average flow rate of 85­369 m s- 1. Although this micropumping principle can be used to perform the continuous-flow PCR with cycle number flexibility, it often suffers from electrolysis that is associated with the formation of bubbles and a drop in the channel current. Especially when operating at PCR temperature conditions electrolysis occurs at considerately lower voltages, resulting in the block of microchannels by bubbles and the formation of electrolysis by-products that could adversely affect the PCR chemistry. In addition, within the MHD micropump, the scaling of flow rate with the fourth power of hydraulic diameter makes miniaturization challenging (Laser and Santiago, 2004). 3.2.3. Electrochemical micropumps Electrochemically generated bubbles in a microchannel can be used as an actuation force. The micropump using this actuation principle is referred to as the electrochemical one. Strictly speaking, it should be called as the nonmechanical `bubble' micropump, where interfacial tension effects take the place of traditional moving surfaces for applying pressure on the working fluid. Recently, Liu et al. used the electrochemical micropumps to drive the sample fluid volumes on the order of mL in the self-contained, fully integrated PCR microchip, as shown in Fig. 15 (Liu et al., 2004b). The micropump relied on electrolysis of water between two platinum electrodes in a saline solution to generate gases when a dc current was applied. The gas generated a pressure that in turn moved liquid solutions inside the chip (Liu et al., 2004b). For pumping of approximately milliliter solution volumes, this micropump is more efficient and consumes less power. For example, a steady flow rate of up to 800 L min- 1 was obtained with a power consumption of b 150 mW (Liu et al., 2004b). 3.2.4. Acoustic-wave micropumps The acoustic-wave can provide an alternative actuation approach for driving the sample fluid. More recently, Guttenberg et al. have reported a planar chip device for PCR and hybridization integrated with surface acoustic wave (SAW) micropump (Guttenberg et al., 2005). The developed PCR microfluidic chip was operated at a planar surface instead of a closed channel network, which can be referred to as a virtual reaction chamber (VRC) based PCR chip. The PCR sample fluid was transported in

Fig. 16. Diagrammatic principle of electrokinetic synchronized cyclic continuous flow PCR process. (A) Sample injection; (B) Sample cycling. Source: adapted from Chen et al. (2005a). [Reprinted with permission from Chen et al. (2005a). Copyright (2005) American Chemical Society.]

the potential and modify electroosmotic flow (Gui and Ren, 2006). 3.2.2. MHD micropumps Microfluid driving is also realized by using a magnetohydrodynamic (MHD) micropump, in which a transversal ionic current inside the microchannel is subjected to a magnetic field oriented in an angle of 90° to current direction and microchannel axis; The Lorentz force acting onto the ionic current in the aqueous solution will then induce a fluid flow in the microchannel direction (Lemoff and Lee, 2000). Using the MHD actuation, West et al. developed a continuous-flow PCR chip integrated

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single droplets using SAW on a piezoelectric LiNbO3 substrate chip that has eight separately addressable SAW transducers. The agitation principle of the SAW micropump can be simply described as follows: the SAWs are produced by interdigitated transducers (IDT) being connected to a high radio frequency (RF) power source (150 MHz). The fast alternating electric field generates displacements on the surface of the piezo with an amplitude in the nm range, moving at speed of sound on the substrate. When a sample fluid drop is placed in the propagation path of the SAW a momentum is transferred to it. Small droplets are actuated in the same direction as the wave when sufficient RF power is applied. This type of micropump based on the effect of acoustic streaming has several obvious advantages: (i) beyond the RF-power threshold increasing the RF power leads to higher speed of the droplets (for example N 5 cm s- 1), allowing for the fast PCR analysis; (ii) the sample volume actuated by the SAW can be lowered up to several picoliter, thus allowing one to perform the picoliter PCR in a chip; (iii) due to the fact that the amplitude of the SAW is decreased rapidly when it reaches the contact line of the liquid and is attenuated by orders of magnitude after several wavelengths, it is very easily used to mix or dispense small droplets, making for the development of PCR chip with multiple analytical functionalities; (iv) finally, the SAW actuation depends only slightly on the components of the sample fluid (for example ion density and pH, etc.), compared with the electrokinetic or MHD actuation. 3.2.5. Surface tension and capillary micropumps Capillary forces can be controlled actively or passively using different effects: thermocapillary, electrocapillary or passive capillary. The thermocapillary effect is caused by the temperature gradient on the surface. The electrocapillary effect, also known as electro-wetting, is induced by the localized potential difference on the surface. The use of passive capillary effect depends on the geometries or the surface properties (hydrophobic or hydrophilic surfaces). Here, we will in brief introduce the applications of the capillary-effect-based micropumps to the integrated PCR chips. Early in 1996, Burns et al. integrated the thermocapillary micropump with a DNA analysis microchip that consisted of thermal-cycling chambers, gel electrophoresis channels, and radiolabeled DNA detectors (Burns et al., 1996). The thermocapillary micropump provided movement of discrete sample liquid drops in microchannels without moving parts or valves, only by differentially heating the drop inter-faces. In microchannels, a pressure difference, which is a function of the surface tension, occurs across the liquid-air interface (i.e., capillary

pressure). By using this thermocapillary micropump, the moving, mixing, and splitting of small-volume sample liquid drops can be very easily carried out. For example, the sample drops of 60 nL were driven and mixed in the Y-shaped microchannel by thermocapillary pumping (Burns et al., 1996). In addition, noteworthy is that a pumping system based on individual drop movement can provide two additional advantages for DNA analysis: (i) the sample volume can be determined by measuring the drop length, and (ii) each sample is kept separate, decreasing the risk of cross-contamination. As a second capillary micropump concept, electrocapillary actuation has been developed for microfluidic PCR applications (Pollack et al., 2003; Chang et al., 2006). Within the electro-wetting-on-dielectric (EWOD) chip, Pollack et al. performed the real-time PCR assays in 300 nL droplets containing standard PCR reagents. The manipulations and the environment inside the electrowetting chip did not inhibit PCR (Pollack et al., 2003). However, this system used an additional PCR device including an external thermocouple and heating fans for PCR temperature control. In a similar approach, most recently, Chang et al. reported an integrated microfluidic chip for PCR applications utilizing digital microfluidic chip (DMC) technology, in which sample transportation, mixing, and DNA amplification were accomplished by EWOD effect (Chang et al., 2006). With the help of the passive capillary effect (i.e., the hydrophobic (Teflon®)/ hydrophilic (SiO2) structure), the DMC chip was robustly integrated with the PCR chip device. Sample droplets were first generated, driven and mixed by the EWOD actuation. Then the mixture droplets (730 nL) were transported to the PCR chamber integrated with two film heaters and a film temperature sensor by using the hydrophobic/ hydrophilic interface to generate required surface tension gradient, and this driving process took 0.13 s. The integrated DMC/PCR chips only required an operation voltage of 12 VRMS at a frequency of 3 KHz for digital microfluidic actuation and 9 VDC for thermal cycling. 3.2.6. Ferrofluidic magnetic micropumps Ferrofluids are magnetic liquids created by suspending ferromagnetic particles in a carrier fluid. Carrier fluids can be water, hydrocarbons or fluorocarbons and favor many different applications. Ferrofluids conform to the channel shape, potentially providing very good seals, and respond to external localized magnetic forces, providing easy actuation. The use of ferrofluids as micropumps for simple and fast nucleic acid amplification in the integrated chip was presented by Münchow et al. (Münchow et al., 2005). The ferrofluid used in this device was a suspension of ferromagnetic nanoparticles in an oil carrier. The ferrofluid

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as a piston was moved forward and backward by a magnet that was connected to a stepper motor and could be moved above the ferrofluid channel, as shown in Fig. 17. The maximum pressures achievable with the ferrofluidic pump were of the order of 0.1 kPa, but this value increases with a decreasing aspect ratio and a decreasing hydraulic diameter of the microchannel. In addition, the presented ferrofluidic magnetic micropump can provide a very precise, pulsationfree and reliable sample fluid driving, and the deviation between the first and the last (40th cycle) position of the sample plug was less than 100 m. Also, since the ferrofluid plug was only connected to the PCR sample via pneumatic transduction, the risk of cross-contamination was minimized. It is noted that the ferrofluidic micropumps can be not only used for sample plug transport between the PCR different temperature zones, but also for the metering and merging process (Münchow et al., 2005). 4. Fluid blending: micromixer for PCR microfluidic chips Besides the microvalve and micropump, the micromixer is often another important and indispensable component in an integrated PCR microfluidic system. Early in the development of microfluidic PCR chips, the importance of micromixers within the PCR chips was not well recognized and only a few groups were focused on this research field (Burns et al., 1996, 1998; Mastrangelo et al., 1999; Anderson et al., 2000). Recently, many new micromixers have been desirably used inside the integrated PCR chips (Chou et al., 2001; West et al., 2002; Liu et al., 2003a; Curcio and Roeraade, 2003; Liu et al., 2004b; Lee et al., 2005; Dorfman et al., 2005; Hashimoto et al., 2006; Chang et al., 2006; Chabert et al., 2006; Lien et al., 2007; Mohr et al., 2007). In fact, most of the works reviewed in this section were published in the past 5 years. In general, two basic principles are followed

to produce mixing at the microscale: passive micromixing and active micromixing. Passive micromixers do not require external energy, and the mixing process depends only on diffusion or chaotic advection. For this micromixing means, the flow energy due to pumping action or hydrostatic potential is used to restructure a flow in a way which causes faster mixing. Active micromixers use energy input from the exterior for the mixing process. These external energy sources include electrokinetics, acoustics, magnetohydrodynamics, integrated microvalves/micropumps, and others. 4.1. Passive micromixers for PCR chips 4.1.1. Y/T-type flow micromixers The Y/T-type flow configurations are simple mixing structures, but have been well applied for sample fluid mixing in the PCR chips (Burns et al., 1998; Mastrangelo et al., 1999; Hashimoto et al., 2006; Legendre et al., 2006; Münchow et al., 2005). Burns' group developed a fully integrated DNA chip capable of measuring and mixing reagent and sample droplets, amplifying or digesting the DNA, and finally separating and detecting products (Burns et al., 1998; Mastrangelo et al., 1999). Within this chip, the reagent and sample drops introduced in the injection region were first mixed by four T-type micromixers, they were then further mixed by a Y-type micromixers in the mixing region. Although the structure of this T/Yintegrated micromixer is complicated and requires complex fabrication processes, it can considerably increase the mixing efficiency. A single Y-shaped passive micromixer was recently used to mix the continuous-flow PCR product and ligation detection reaction (LDR) cocktail for the detection of low-abundant DNA point mutations (Hashimoto et al., 2006). The PCR, LDR and micromixing were performed in a PC chip. Although this work did not detail the performance of this Y-shaped micromixer, the mixing efficiency should be very desirable, as suggested from the experimental results. Münchow et al. also first utilized a simple passive Y-type microchannel to induce the reproducible metering and merging of two or more fluid volumes (0.1­10 L) without any active valve elements or external sensors, and then the species inside the merged sample plug were mixed by increasing the fluid interface between them due to a recirculation flow inside of the moving droplets (discussed below). During this micromixing process, the plug was transported into the PCR chip for actual amplification (Münchow et al., 2005). Most recently, Legendre used a simple T-shaped flow configuration to combine the lateral diffusion mixing and the thermally-induced convective mixing (Legendre et al., 2006). Diffusive mixing occurred during transfer from the

Fig. 17. Ferrofluidic magnetic micropump scheme for the oscillating-flow PCR chip. Source: Münchow et al. (2005). [Reprinted with permission from Münchow et al. (2005). Copyright (2005) Future Drugs Ltd.]

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solid-phase extraction (SPE) bed to the PCR chamber and then in the chamber itself, while convective mixing resulted from thermocycling during PCR. By this coupled micromixing strategy, the eluted DNA from the SPE domain and the PCR master mix from the side arm were completely mixed in 2 min. However, the diffuse micromixing of the solutions without any heating showed incomplete even after 10 min. 4.1.2. Recirculation flow micromixers Anderson et al. first utilized either porous hydrophobic vents alone or in combination with valves to joint or link reagents, and the resultant reagents were then mixed by the recirculation flow micromixing strategy so as to achieve homogeneity (Anderson et al., 2000). In general, mixing fluids in small channels with low Reynolds number (laminar low) is considered difficult, but propelling a bolus or packet of fluid through a channel creates an advantageous recirculation pattern along the axis of the channel. This is a consequence of non-slip boundary conditions, superimposed with the requirement that the leading and trailing menisci move as one. The recirculation effect facilitates effective blending of elongated aqueous reagent packets. 4.1.3. Droplet micromixers In order to reduce the mixing path to increase the mixing rate and efficiency, an alternative solution is to form droplets of the mixed sample liquids. The moving of a droplet produces an internal flow field and makes mixing inside the droplet possible. In general, sample droplets can be produced and transported individually by using capillary effects such as thermocapillary (Burns et al., 1996) and electrocapillary (i.e., electro-wetting) (Chang et al., 2006). Moreover, sample droplets can be generated due to the large difference of surface tensions in a small channel with multiple immiscible phases such as sample aqueous solution/Perfluorodecalin organic liquid (Curcio and Roeraade, 2003), sample aqueous solution/FC-40 fluorinated oil containing the fluoro alcohol surfactant (Dorfman et al., 2005; Chabert et al., 2006), and sample aqueous solution/ light mineral oil or sunflower oil (Mohr et al., 2007). As mentioned early, sample droplets could be moved via thermocapillary, where the electro-thermal control of interfacial tension was proposed as a microscale pumping mechanism (Burns et al., 1996). If sample droplets (60 nL) were loaded into the two channels of the Yshaped micromixer and were driven to the channel intersection by thermocapillary, they were merged to form a single large droplet; the combined droplets could be stopped by turning off all microheaters and could be reversed by heating the right interface (Burns et al., 1996).

This mixing is achieved by droplet movement only, and therefore it should be passive mixing owing to convections. It needs to be noted, additionally, that circulation patterns generated inside the droplet during motion help to mix the liquid sample. By simplifying the mass transport equation and introducing an effective dispersion coefficient for a rectangular channel, Burn's group also reported an analytical model for droplet mixing driven by thermocapillary (Handique and Burns, 2001). The droplet micromixer can also be moved by electrowetting which is the change of surface energies by applying electric fields. More recently, Chang et al. reported a mixing scheme with the EWOD concept for the integrated PCR chip (Chang et al., 2006). The two droplets containing PCR reagents and cDNA samples were first merged to form a bigger droplet. Then the merged droplet was moved counterclockwise around a square loop consisting of only the 2 ×2 electrode array. By switching the applied voltages of the electrodes in specific loops, the merged droplets could be well mixed after three cycles of the counterclockwise movement. The mixing efficiency was 20.5% when two droplets were initially merged, while the mixing efficiency of the merged droplet was measured to 92.8% after 3 cycles of the counterclockwise motion in the specific loops (Chang et al., 2006). The other droplet micromixer architecture used flow instability between two immiscible liquids (Curcio and Roeraade, 2003; Dorfman et al., 2005; Chabert et al., 2006; Mohr et al., 2007). Using a carrier liquid such as Perfluorodecalin (Curcio and Roeraade, 2003), fluorinated oil FC-40 containing fluoro alcohol surfactant (Dorfman et al., 2005; Chabert et al., 2006), and light mineral oil or sunflower oil (Mohr et al., 2007), droplets of the PCR aqueous samples were formed. While moving through the microchannel, the shear force between the carrier liquid and the sample accelerated the mixing process in the droplet. The use of surfactant can prevent the transient adsorption of droplets to the microchannel walls. 4.2. Active micromixers for PCR chips 4.2.1. Electrokinetically-driven micromixers As mentioned above, electrokinetic flow can be used to transport sample fluid as an alternative to pressure-driven flow. In a different light, the electrokinetic disturbance (or instability) can also be used as the micromixing mechanism (Jacobson et al., 1999), where mixing is accomplished by the action of fluctuating electric fields. Recently, to enhance the homogeneous mixing of samples and reagents in the PCR chamber, Lee et al. developed an active electrokinetically driven micromixer using zeta potential variations

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in the silicon-based microchannels (Lee et al., 2005). Variation of zeta potential on the wall of microchannels was induced by using five pairs of buried shielding electrodes, rather than using only one pair of electrodes, which can shorten the fully mixing length within 4500 m. This work showed that without the activation of the micromixers the fluid lacked the transverse transformation due to the smooth streamlines and consequently the mixing in the microchannel was inefficient; by alternating the shielding electrodes with a given frequency at a certain control voltage, the local circulation zones in the microchannels were formed, and moreover the buck flow followed a significantly narrower region and more complicated route, therefore increasing the rate of diffusion by the local variation of zeta potential. In addition, it is noted that the mechanism of the stretching and folding will exist in the fluid flow as if the shielding electrodes are driven with a given frequency. As a very important criterion for valuating the micromixers, the time constant of this active micromixer was only around 4 s. Note that this time constant can vary due to dimension of the microchannels (Lee et al., 2005). 4.2.2. Acoustically-driven micromixers An air bubble in a liquid medium can act as an actuator (i.e., the bubble surface behaves like a vibrating membrane) when it is activated by a sound field (Liu et al., 2002b, 2003b). The behavior of bubbles in sound fields is determined largely by their resonance characteristics. Bubble vibration within a sound field induces frictional forces at the air/liquid interface which produce a bulk fluid flow around the air bubble, called cavitation microstreaming or acoustic microstreaming. The bubble-induced streaming strongly depends on frequency for a given bubble radius, and on bubble radius for a given frequency. Acoustic microstreaming arising around a single bubble excited close to resonance causes strong liquid circulation flow in the liquid chamber, which can be used to effectively enhance mixing. The micromixing technique based on acoustic microstreaming principle has been developed to accelerate DNA hybridization process or to increase the immunomagnetic cell capture process efficiency (Liu et al., 2002b, 2003b, 2004b). Recently, these processes have been incorporated onto a self-contained, full integrated biochip for sample preparation, PCR amplification, and DNA microarray detection, as shown in Fig. 15 (Liu et al., 2004b). In this micromixer, air pockets with a 500 m diameter and 500 m in depth were used for trapping air bubbles. Acoustic microstreaming was induced by the field generated by an integrated PZT actuator. Microfluidic experiments showed that the time taken to achieve a complete micromixing in a 50 L chamber using cavitation microstreaming was reduced from hours (a pure diffusion-

based mixing) to only 6 s. Cell capture efficiency of 73% was achieved, as compared to 91 % using conventional vortexing in a microfuge tube and 2% using pure diffusion in the same chamber microdevice. In addition, acoustic microstreaming resulted in up to 5-fold hybridization reaction kinetics acceleration with significantly improved signal uniformity. Acoustic microstreaming has many advantages over most existing chamber micromixing techniques, including simple apparatus, ease of implementation, low power consumption (2 mW), and low cost (Liu et al., 2002b, 2003b, 2004b). 4.2.3. MHD-driven micromixers The MHD effect has been used in micromixers (Bau et al., 2001). As mentioned above, in the presence of an external magnetic force, applied DC voltages on the electrodes generate Lorentz forces, which in turn induce mixing movement in the chamber. The lorentz force can roll and fold the electrolyte solutions in a mixing chamber. The MHD micromixer of Bau et al. was fabricated from low temperature co-fired ceramic tapes (LTCC). The electrode arrays, which were printed with a gold paste, were deposited on the conduit's surface in the transverse direction, instead of being positioned parallel to the conduit walls (Bau et al., 2001). By a slightly different approach, West et al. reported a MHD micromixer for continuous-flow PCR amplification, which utilized arrays of electrodes deposited on the walls of a silicon or SU-8-based microchannel (West et al., 2002). Through alternate AC potential differences across pairs of electrodes, currents in various directions of the mixing volume were generated. By variable electrode patterning, complex flow fields could be induced. Therefore, a so-called cellular motion was initiated to enlarge the fluid interfaces for mixing. 4.2.4. Micromixers with integrated microvalves/ micropumps The organic integration of microvalves with micropumps can provide an alternative approach for active micromixing of samples and reagents. For example, Lien et al. most recently reported a micromixer consisting of pneumatic microvalves and micropumps for mixing the Dengue virus samples and the antibodyconjugated magnetic beads, as shown in Fig. 9 (C) (Lien et al., 2007). Viruses were bound onto the magnetic beads in the incubation process and purification was performed in the washing process. By using the on-chip microfluidic mixer, the time taken to perform the incubation and washing processes was decreased from approximately 12 h (a conventional large-scale mixer) to only 4 h (Lien et al., 2007).

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Quake's group utilized a pneumatic peristaltic rotary micropump for two kinds of on-chip mixing: fixedvolume mixing and continuous-flow mixing (Chou et al., 2001). The former is the case in which one can mix two different solutions (sample and reagent) of fixed volume completely before they go on to another stage of processing. Instead of injecting them into a chamber and waiting for the slow diffusing to take place, they can be put into the circulation loop and quickly mixed with the rotary pump. The mechanism behind this is the parabolic profile of the Poiseuille flow. As two fluids rotate in the same loop, the center of each fluid flows faster than the edge. Therefore, the interface between them keeps stretching and eventually each fluid becomes a long and thin stream that wraps around with each other. Continuous-flow mixing is essentially dynamic mixing, which takes place when two solutions are flowing down the channel continuously. Since this mixing is done by feeding a well-mixed solution back into the original flow stream, the quality of the mixing depends strongly on the ratio of the overall flow rate and the pump rate. Better mixing could be achieved by either lowering the overall flow rate or by widening and lengthening the rotary mixer loop. A simple analytical model was developed to estimate the mixing effect of rotary flow, and instead of a time factor of 1000 in passive mixing devices, the time required for mixing was only a factor of 10 between fast and slow objects when rotary pumping was used (Chou et al., 2001). Such a micromixing device has been successfully applied by Quake's group onto a nanoliter continuous-flow PCR chip device (Liu et al., 2002a) or a 400-throughput PCR chip (Liu et al., 2003a). The volumes loaded were 3 L of primers and noprimer control, 1 L of cDNA or no-template control, and 2 L of Dynazyme. An efficient mixing of the samples and/or reagents was achieved by actuating the rotary pumps with two pneumatic controllers at 10 Hz for 5 min (Liu et al., 2003a). 5. Discussion Throughout this review, the development of microvalves, micropumps and micromixers in microfluidic PCR chips has been surveyed with respect to various operation mechanisms, as will be discussed in the following sections. 5.1. Microvalves for microfluidic PCR chips The functions of the microvalves in PCR chips include flow regulation, on/off switching, or sealing of PCR sample fluids. They are critical to successful integration of PCR amplification with other DNA assays such as

hybridization and CE separation. Such microvalves have to meet a number of requirements. First, the valves have to be able to withhold the pressure generated during PCR, caused by degassing and air expansion at elevated temperature, to ensure the successful confinement of the PCR sample inside the PCR chamber during PCR. Second, after PCR, the valve should be easily opened so as to allow PCR solution to flow into the assay chamber/ channel. Third, because valves will be in direct contact with PCR solution, the valve material should not inhibit PCR. Finally, other desired characteristics of the microvalves used in the PCR chips include small/zero dead volume, low power consumption, rapid response time, amenability for integration into a microdevice, reproducible valve operation and disposability. Based on these requirements, different approaches have been explored in the development of microvalves for PCR chips. Active mechanical microvalves such as thermallyactivated, normally-open pneumatic microvalves (Baechi et al., 2001) and SMA microvalves (Münchow et al., 2005) have been used within PCR chips. Thermal actuation can provide large forces via large strokes, but is relatively slow during the valve's closing and opening. In addition, the thermally actuated microvalves are easily integrated with other microfluidic functionalities on a single chip, with compact footprint and low operation costs. SMA actuation is preferred in terms of power consumption, since it requires power only during the transition between two stable modes. However, it is difficult to control the displacement precisely in SWA microvalves, so they can only be used as on-off valves. Active non-mechanical microvalves based on phase change actuation have been desirably developed for the integrated PCR chips. The phase change materials used include hydrogel (Wang et al., 2005, 2006), sol­gel (Liu et al., 2002c), paraffin (Pal et al., 2004, 2005; Liu et al., 2004a,b) and ice (Gui and Liu, 2004; He et al., 2001; Chen et al., 2005b). These phase change microvalves based on the solid-liquid transition used thermal actuation of a meltable piston to perform the valve operation without any membrane, and therefore the fabrication and operation of the valve is simple. Most probably, they are very useful in disposable biochip applications due to their relatively low cost. These microvalves can also be electronically addressable and effectively leak-proof, making them amenable for LSI. In addition, the valves can be latched, minimizing the power required and enabling easy fluidic control. However, the time response of these valves is relatively slow as compared to most conventional microvalves ( ms). Table 1 lists the characteristics of thermally-actuated phase change microvalves used in PCR chips.

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Table 1 The characteristics of thermally-actuated phase change microvalves used in PCR chips Reference Type R/I Phase change material Valving channel Mode Fluid Time (s) Pressure Power (kPa) (W) 200 138 1725 External system TE unit TE unit b0.015 Pneumatic pressure/ vacuum

(Wang et al., Hydrogel R 2005, 2006) (Liu et al., Sol­gel I 2002c) (Pal et al., Paraffin R 2004, 2005) (Liu et al., 2004a,b) (Chen et al., 2005b) Paraffin Ice I R

Temperature sensitive hydrogel In-line PC NC (Tc = 32 °C, swelling at T b Tc) Pluronics sol­gel (Tc = 5 °C, liquid at In-line PC NC T b Tc) In-line B M1595 wax (Tm = ~85 °C, phase change) glass/silicon Paraffin (Tm = 70 °C, phase change) Water (phase change) In-line PC In-line PC B NO

PCR 6 (O) mixture 5(C) PCR mixture PCR 0.2 mixture

PCR 10 275 mixture PCR 0.5­10 360 mixture

1.8-6.9 TE unit

Note: R -- reversible; I -- irreversible; NC -- normally closed microvalves; NO -- normally open microvalves; B -- bistable; O -- opening; C -- closing.

Compared with active mechanic and non-mechanic microvalves, active microvalves with external systems have been most widely utilized in the PCR chips due to their excellent performance in on/off switching or sealing (Oh et al., 2005a,b; Anderson et al., 2000; Lagally et al., 2000, 2001a,b, 2003, 2004; Toriello et al., 2006; Liu et al., 2006, 2007, 2002a, 2003a; Hou et al., 2007; Easley et al., 2006; Liao et al., 2005; Lien et al., 2007; Huang et al., 2006; Yuen et al., 2000; Chou et al., 2001; Pilarski et al., 2005; Prakash et al., 2005; Kaigala et al., 2006). Up to now, the pinchtype microvalves with external pneumatic or nonpneumatic actuation forces by indirect contact with any flexible polymer-based membrane and in-line channels (see above) incorporated onto the integrated PCR chips have been favor because they can provide zero leakage flow and large resistible pressure, eliminating the risk of cross-contamination. These microvalves often require an external pneumatic supply and one or more high-speed valve connections and are therefore not strictly compared to microvalves with fully integrated actuators. In settings where the necessary infrastructure is available, however, active microvalves with external systems can be effective. Here, it is strongly recommended that the external systems that provide the external forces to the pinch microvalves should be further miniaturized for portable biochemical applications. Some passive mechanical or non-mechanical microvalves are also incorporated onto the PCR chips to meet integration requirements. Two notable examples are inline polymerized gel microvalves (Fan et al., 2003; Koh et al., 2003) and PP microvalves (Prakash and Kaler, 2007). These microvalves can provide several obvious advantages such as simple fabrication, easy operation, perfect sealing and ability to effectively preventing bulk

flow of liquid during the PCR thermal cycling. Most importantly, these microvalves can allow the PCR chip to be highly integrated and to have a compact footprint. In addition, passive microvalves using capillary effects are often useful for microfluidic PCR applications since autonomous and spontaneous valving can be easily realized due to the surface properties of the microchannels (Gong et al., 2006; Münchow et al., 2005; Liu et al., 2002c; Lagally et al., 2000, 2001a,b; Anderson et al., 2000; Burns et al., 1998). However, the hydrophobic/hydrophilic surface stability associated with these hydrophobic patch microvalves is a problematic issue. Despite this drawback, these passive hydrophobic valves are considered to be useful for blocking and passing fluidic flows without sealing at PCR high temperatures. As a whole, active mechanical microvalves have an important disadvantage of relatively high cost due to their complicated structures and therefore they have not been widely developed for microfluidic PCR chip applications. On the contrary, for life sciences applications such as PCR chips, active non-mechanical and passive microvalves are preferred due to the possibilities of low cost and easy integration into the chip devices, as well as miniaturization of instruments. If the size of the instruments does not matter, microvalves with external systems are desirable. 5.2. Micropumps for microfluidic PCR chips As seem from the above survey, a number of researchers have sought to develop micropumps for transport of a wide range of sample volumes (L­mL) in highly miniaturized PCR chips. On the one hand, the micropumps are used to non-manually drive the PCR

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samples and reagents into the reaction chamber (or channel) from the inlets, and/or the resultant PCR products are expelled from the reaction region for further detection. On the other hand, the micropumps are also used for the development of continuous-flow PCR chips, where the PCR mixture is driven to flow repetitively through three or two different thermostable zones (Chen et al., 2005a; Hu et al., 2006; Gui and Ren, 2006; West et al., 2002; Münchow et al., 2005). In addition, the proper combination of micropumps with microvalves can be utilized for active micromixing of PCR samples and reagents (Lien et al., 2007; Liu et al., 2002a, 2003a). The desired characteristics of the micropumps include compact device size, easy integration into a chip, stability and reproducibility of fluid flow, compatibility with samples and reagents, insensitivity to particulate contamination, large pump pressure, convenient management, long service life, ability to operate with various fluids and disposability. To meet these requirements, various approaches have been explored in the development of micropumps for PCR chips. Of course, there is no such micropump that meets all requirements dimensioned above. Several kinds of mechanical micropumps have been desirably used in the PCR chips, including piezoelectric (Bu et al., 2003), pneumatical (Liao et al., 2005; Lien et al., 2007; Huang et al., 2006; Chou et al., 2001; Liu et al., 2002a, 2003a; Easley et al., 2006), thermopneumatic (Liu et al., 2004b; Burns et al., 1998; Mastrangelo et al., 1999), and electric­servomotor-driven (Pilarski et al., 2005; Prakash et al., 2005; Kaigala et al., 2006) micropumps. These micropumps have two movement

Table 2 Mechanical micropumps reported in PCR chips Reference Bu et al. (2003) Liao et al. (2005) Lien et al. (2007) Huang et al. (2006) (Chou et al., 2001; Liu et al., 2002a, 2003a) Easley et al. (2006) Driver Check valves Construction Glass-silicon PDMS-glass PDMS-PDMSglass PDMS-PDMSglass-PMMA Multi-layer elastomer (PDMS)

mechanisms: diaphragm (Easley et al., 2006) and peristaltic (Bu et al., 2003; Liao et al., 2005; Lien et al., 2007; Huang et al., 2006; Chou et al., 2001; Liu et al., 2002a, 2003a; Pilarski et al., 2005; Prakash et al., 2005; Kaigala et al., 2006) movements. For diaphragm micropumps, the basic components are a pump chamber, an actuator mechanism or driver and two passive check valves--one at the inlet (or suction side) and one at the outlet (or discharge side), and the performance of check microvalves at the inlet and outlet of the pump chamber is critical to the operation of diaphragm. Micropumps with multiple chambers in series and no valves, or micropumps with multiple active microvalves operated in a series of on/off actuation sequences are sometimes referred to as peristaltic micropumps. In the PCR chips, the peristaltic micropumps are preferred probably due to its simplicity in design and operation. In addition, a peristaltic micropump can allow bidirectional or unidirectional sample fluid transport by simply changing the active microvalve actuation scheme. However, it should be noted that performance improvements realized with a multi-chamber or multi-active-valve design in the peristaltic micropumps must be balanced against increases in fabrication complexity and overall device size. Table 2 summarizes some key features and measured performance characteristics of diaphragm and peristaltic micropumps reported in the PCR chips. As with mechanical micropumps, non-mechanical dynamic micropumps have been emerging in the research field of PCR microfluidic chip, including electrokinetic (Lee et al., 2005; Chen et al., 2005a; Hu et al., 2006; Gui and Ren, 2006), MHD (West et al., 2002), electrochemical

Pump Diaphragm Diaphragm V f pmax Qmax chamber material thickness (mm) (V) (Hz) (kPa) (mL s- 1) 3 (S) 3(S) 3(S) 3(S) 6 (S) Glass PDMS PDMS PDMS Elastomer PDMS PDMS PDMS 0.254 1 12 5­7 60 0.0644 0.2 100 10 9 15 11 50-60 69 173 3.14 0.0667 1.67 0.1

Piezoelectric None (axial) Pneumatic None Pneumatic None Pneumatic Pneumatic None None

0.3

15 138

Pneumatic

(Pilarski et al., 2005; ElectricPrakash et al., 2005; servomotor Kaigala et al., 2006)

Flap Glass-PDMS(diaphragm) glass-glass None PDMS-glass

1 3(S)

Note: S -- series configuration; V -- micropump operating voltage; f -- micropump operating frequency; pmax -- maximum measured micropump differential pressure; Qmax -- maximum measured volumetric flow rate.

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(Liu et al., 2004b), acoustic-wave (Guttenberg et al., 2005), surface tension and capillary (Burns et al., 1996; Pollack et al., 2003; Chang et al., 2006), and ferrofluidic magnetic (Münchow et al., 2005) micropumps. Among these nonmechanical micropumps, electrokinetic micropumps are preferred due to the possibilities of low cost, simply system design and easy integration into the LOC analytical devices, as well as high miniaturization of instruments. Another important advantage of electrokinetic flow micropump is that fluid flows in a microfluidic network can be controlled easily by switching on/off voltages; this control circumvents the need for valves. However, electrokinetic flow micropump has several important drawbacks, including buffer incompatibility, frequent changes of voltage settings, electrolytic bubble formation, and evaporation of solvent due to heating. Also, electrophoretic demixing is disadvantageous in assays requiring a uniform flow for all species. For MHD micropump, it provides several desired characteristics, including easy architecture and fabrication, no pulsation of fluid flow and bidirectional adjustability of fluid flow. Nevertheless, when a MHD micropump is operated, the electrolytic bubbles very easily cause the block of microfluidic channels and the formation of electrolysis by-products will also inhibit the PCR chemistry. The electrochemical micropump is actually a "bubble" one where the bubbles are electrochemically produced, and therefore its working principle is sometimes similar to that of thermopneumatic micropump with no diaphragm membrane actuation (Liu et al., 2004b). This micropump method is very simple, compared with conventional pressure-driven membrane-actuated micropump ones that often suffer from complicated designs, complicated fabrication, or high cost. Capillary micropumps are usually based on interfacial tension resulted from electrocapillary, thermocapillay or hydrophobic/ hydrophilic effects. Although the operation principle associated with these micropumps is simple, it is not very convenient to actively alter liquid surface tension inside microchannels so as to provide continuous and controllable driving forces. The micropumps based on SAW actuation can provide an alternative approach for driving the sample fluid on the LOC device and have obvious superiorities over other micropumps (see above). However, the construction process of SAW micropump is complicated and the cost of the whole chip is considerably high. Moreover, in such micropumps, the biocompatibility of chip substrates such as LiNbO3 that are not routinely used in biochip field, needs to be further investigated. In a word, non-mechanical dynamic micropumps based on electromagnetic, acoustic and capillary effects are a subject of increasing interest, and their superiorities have not yet brought into use in PCR chips. However, it

is the author's belief that these non-mechanical micropumps, especially electrokinetic micropumps are emerging as a viable option for fully integrated PCR chips. 5.3. Micromixers for microfluidic PCR chips PCR is a multi-component biochemical reaction and the components include master mix (magnesium salts, pH buffer, and potassium), DNA template(s), oligonucleotide primers, polymerase enzyme, and deoxyribonucleotide triphosphates (dNTPs). Therefore, before a real PCR amplification, the PCR solution should be effectively mixed so as to obtain the good reaction result. However, to date, most of such operations have been manually accomplished by the conventional vortexing in a microfuge tube. Apparently, such operation processes can't meet the requirement of TAS assay. In order to circumvent this issue, the passive and active micromixers have been developed for non-manually mixing the PCR samples and reagents on a monolithic chip. Mixing in passive micromixers relies mainly on molecular diffusion or chaotic advection due to the dominating laminar flow on the microscale. To passively accomplish effective micromixing of PCR solutions, various passive micromixers have been reported, including Y/T-shaped (Burns et al., 1998; Mastrangelo et al., 1999; Hashimoto et al., 2006; Legendre et al., 2006; Münchow et al., 2005), recirculation flow (Anderson et al., 2000) and droplet (Burns et al., 1996; Chang et al., 2006; Curcio and Roeraade, 2003; Dorfman et al., 2005; Chabert et al., 2006; Mohr et al., 2007) micromixers. Since the basic T/Y-mixer entirely depends on molecular diffusion, a long mixing channel is usually needed, which isn't in favor of the development of integrated PCR chip with compact footprint. To overcome this problem, a couple mixing model with short-channel T-type micromixing and in-chamber thermally-convective micromixing has been reported (Legendre et al., 2006). The dropletbased micromixing is a very promising mixing method. On the one hand, because of the reduction of mixing path inside the droplet micromixer, the effective and rapid mixing can be expected. On the other hand, the dropletbased actuation can provide an alternative approach for realizing the valveless, digital microfluidic PCR platforms without loss of desired analytical functionalities (Guttenberg et al., 2005; Burns et al., 1996; Chang et al., 2006). The droplets can be created, driven, divided or mixed by the electro-wetting (Chang et al., 2006), thermocapillary (Burns et al., 1996), or SAW (Guttenberg et al., 2005) actuation. In addition, it should be noted that the electrowetting effect can also induce an active micromixing by shaking the merged droplet. This turned out to be

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particularly effective when using multi-electrode arrasys (Chang et al., 2006). Different from passive micromixing, active micromixers use an external field to produce the disturbance for the mixing process. Thus, almost all external disturbance effects can be applied to induce the active micromixing, including pressure, temperature, electrohydrodynamics, electrokinetics, magnetohydrodynamics, acoustics, dielectrophoretics, and so on. Up to now, several kinds of active micromixers have been developed for the integrated PCR chips, for example electrokinetic (Lee et al., 2005), acoustic microstreaming (Liu et al., 2004b), MHD (West et al., 2002), electro-wetting (Chang et al., 2006), and pressuredriven microvalves/pumps-based micromixers (Lien et al., 2007; Liu et al., 2002a, 2003a). Among these active micromixers, the electrokinetic or integrated microvalves/ micropumps-induced micromixer may be promising due to the fact that the microvalve/micropump/micromixer can be accomplished by a single actuation principle (for example pneumatic or electrokinetic actuation) on a monolithic PCR chip, which reduces problems in utilizing multiple actuation approaches. As is mentioned above, micromixing induced by MHD actuation very easily leads to the formation of bubbles that makes against the PCR chemistry. The acoustic microstreaming mixer is based on the bubble vibration within a sound field, where an air bubble in a liquid medium acts as an actuator. Although this micromixing scheme has been successfully utilized to achieve homogeneous mixing of magnetic beads and target cells from the blood suspension (Liu et al., 2002b) and to enhance microarray hybridization detection of the PCR amplicons (Liu et al., 2003b, 2004b), it is challenging in mixing the PCR solutions inside (the upstream channel of) the PCR chamber due to the non-compatibility between air bubble and PCR chemistry. It needs to be noted, in addition, that with the external fields and the corresponding integrated components, the structures of these active micromixers are usually complicated and require complex fabrication processes. Moreover, external power supplies are needed for the operation of active micromixers. Therefore, the integration of active micromixers in a PCR chip is still a challenging task. 6. Conclusions The survey given in this paper reveals remarkable progress in the development of microvalves, micropumps, and micromixers within integrated PCR chips. Various kinds of microvalves, micropumps and micromixers with different designs and structures have been reported within PCR chips. The performance of these microfluidic components has been constantly improved and features (for

example leakage, resistible pressure, dead volume, response time, biochemical compatibility of microvalves; compactness, stability, reproducibility, pump pressure and rate, service life of micropumps; mixing speed, mixing efficiency, biocompatibility, compactness, power consumption of micromixers) have been partially addressed and solved. However, the microvalves and micropumps so far reported in the literature for integrated PCR-chip production still have their respective major shortcomings. For microvalves, mechanical microvalves have relatively high cost due to their complicated designs and fabrication processes, and therefore are difficult to meet the requirements for disposable PCR chip applications; active external microvalves mostly used within PCR chips often need a bulky external pressure supply and high-speed valve connections that are incompatible with the PCR chip with a small footprint with respect to their volumes. As for micropumps, especially for pneumatical micropumps, their footprints and compact indexes are relatively large, and therefore are unfavorable for the development of compact microfluidic PCR chip devices. In addition, the compatibility of non-mechanical, electrokinetic or MHD micropumps with microfluidic PCR amplification reaction is also a challenging issue (see Sections 3.2.1 and 3.2.2). The potential breakthroughs in the designs and development of these microfluidic elements within PCR chips are included as follows. First, as biocompatible and inexpensive polymer materials are increasingly being developed and the MEMS fabrication techniques are increasingly evolved, the low-cost, ease-to-use disposable microvalve/micropump/micromixer elements can be incorporated into the highly-integrated PCR chips. Second, the development of active, "self-supplied" compact microvalves/micropumps within PCR chips would be an important research area, where no external bulky pneumatic supply, valve connection tubing, and even power supply are needed. Third, electrically-controlled microvalves/micropumps/micromixers coupled with precise droplet manipulation (Link et al., 2006; Jensen and Lee, 2004) within PCR chips will be emphasized. Fourthly, the integration scale of microvalves/micropumps/micromixers on a single chip (Felton, 2003) will increasingly get larger so as to perform the high-throughput PCR amplification and subsequent analysis. In short, there is plenty of room for improving the performance of these microfluidics components and making them cost-effective for further commercialization. Also, noteworthy is that a number of existing microvalving, micropumping, and micromixing techniques (Oh and Ahn, 2006; Laser and Santiago, 2004; Woias, 2005; Nguyen and Wu, 2005; Hessel et al., 2005; Thielicke and Obermeier, 2000) have not yet been applied towards improving integrated PCR microfluidic chips.

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Nevertheless, as the reliability and ease of manufacture of microvalves, micropumps and micromixers improves, we can expect that these three important microfluidic components will be increasingly used in the PCR microfluidic chip field and other fields such as semiconductors and space life exploration. There is no doubt that in the future that PCR microfluidic platforms embedded with microvalves, micropumps and/or micromixers will be as prevalent as computers are today. Acknowledgements We acknowledge the financial support by the National Natural Science Foundation of China (30670507; 30600128; 30470494) and the Natural Science Foundation of Guangdong Province (015012). The authors would like to express sincere regrets to all researchers who are working on both laboratory research and/or the practical commercialization of micropump, microvalve and/or micromixer devices, while whose work could not be cited in this review due to limited space. References

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