Fully integrated microfluidic separations systems for biochemical analysis

Journal of Chromatography A, 1168 (2007) 170–188

Review

Fully integrated microfluidic separations systems for biochemical analysis Gregory T. Roman a , Robert T. Kennedy a,b,∗ a

b

Department of Chemistry, University of Michigan, USA Department of Pharmacology, University of Michigan, USA Available online 12 June 2007

Abstract Over the past decade a tremendous amount of research has been performed using microfluidic analytical devices to detect over 200 different chemical species. Most of this work has involved substantial integration of fluid manipulation components such as separation channels, valves, and filters. This level of integration has enabled complex sample processing on miniscule sample volumes. Such devices have also demonstrated high throughput, sensitivity, and separation performance. Although the miniaturization of fluidics has been highly valuable, these devices typically rely on conventional ancillary equipment such as power supplies, detection systems, and pumps for operation. This auxiliary equipment prevents the full realization of a “lab-on-a-chip” device with complete portability, autonomous operation, and low cost. Integration and/or miniaturization of ancillary components would dramatically increase the capability and impact of microfluidic separations systems. This review describes recent efforts to incorporate auxiliary equipment either as miniaturized plug-in modules or directly fabricated into the microfluidic device. © 2007 Elsevier B.V. All rights reserved. Keywords: Microfludidics; Capillary electrophoresis; Instrumentation; Microfabrication

Contents 1. 2.

3.

4.

5.

Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Integration of detection transducers onto microfluidic devices . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.1. Fabrication techniques for transducer integration . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.2. ECD integration . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.3. Fluorescence detection integration . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.3.1. Integrated excitation sources . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.3.2. Integrated waveguides and fiber optics . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.3.3. Integrated lenses and filters . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.3.4. Integrated light detectors . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.3.5. Other detectors . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Integration of pumps and power supplies for fluid manipulation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.1. Electrokinetic pumps . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.2. Microfluidic peristaltic membrane and piezoelectric pumps . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.3. Electrical power supplies and control . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Complete systems for biochemical analysis . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.1. DNA and RNA . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.2. Proteins and peptides . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.3. Amino acids, neurotransmitters, and small metabolites . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.4. Applications . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

DOI of original article:10.1016/j.chroma.2007.06.009. Corresponding author at: 930 N. University Ave., Department of Chemistry, University of Michigan, Ann Arbor, MI 48109-1055, USA. Tel.: +1 734 615 4363; fax: +1 734 615 6462. E-mail address: [email protected] (R.T. Kennedy). ∗

0021-9673/$ – see front matter © 2007 Elsevier B.V. All rights reserved. doi:10.1016/j.chroma.2007.06.010

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1. Introduction Over the last decade microfluidic technologies have burgeoned into a multi-million dollar industry with chemical analysis, especially separations, emerging as one of the prime application areas. The enormous interest in microfluidics has been spurred in part by realization that properties at the microscale, such as laminar flow, small diffusion distance, high surface area to volume ratio, and potential for electroosmotic flow can be exploited to yield improved performance for analytical methods [1,2]. High-speed electrophoretic separations for example, are facilitated in micrometer scale conduits [3–5]. Another important factor driving microfluidics development has been with the use of microfabrication methods based on common semiconductor processing techniques. These methods have made it possible to develop complex fluidic circuits that integrate multiple analytical and sample processing functions, such as sampling, filtering, derivatization, dilution, and separation onto a monolithic device [5–11]. This integration of function results in highly automated and rapid measurements even when considerable complexity is required. Because of the small size of the channels used, low volumes (nanoliters and less) are manipulated resulting in reduced reagent consumption and facilitated analysis of microscale samples. Microfabrication also facilitates development of parallel architectures for analytical measurements yielding potential for high-throughput analysis. Automated microfluidic devices have proven capable of performing thousands of chemical assays per day [12,13] and in the near future could potentially perform millions of assays per day. These various benefits can be combined. For example, rapid single cell or cell-group analysis may be performed on microfluidic chips to provide an efficient route to determine statistical variation among biological entities [13,14]. Such methods take advantage of integration of fluidic function, miniaturization of sample size, improved properties for analysis, and high-throughput capability. Because microfluidic devices can be fabricated from a number of inexpensive materials it is also possible to develop disposable devices [15,16]. Microfabrication of fluidic devices also has generated interest because of the potential for high-performance analytical tools with small footprint, low reagent consumption, and modest power requirements. Indeed, the advent of microfluidic analysis resulted in the concept of a “lab-on-a-chip”. This term conjures the idea of a small, completely self-contained, highly portable device with advanced chemical analysis capability. The development of such devices could revolutionize analytical chemistry by making chemical measurements cheaper and more accessible. A truly integrated lab-on-a-chip would have impact in many areas including: (1) diagnostics, e.g. by enabling routine bedside monitoring, assays in the home, or simplified tests in Third World countries; (2) environmental monitoring in industrial settings where many assays must be performed to ensure health and safety of workers or in extreme settings such as in spacecraft; and (3) biomedical research by lowering costs for expensive projects such as genomics, proteomics, or metabolomics measurements. A lab-on-a-chip would also enable remote analysis such as measurements under hazardous conditions.

171

A complete lab-on-a-chip is an attractive idea, but considerable effort is still required to make it a reality for most types of measurements. Although enormous progress has been made in developing fluidic manifolds or networks that are capable of processing samples on chips, in some cases chips smaller than a US dime [17], nearly all chips that have been reported to date require external ancillary equipment such as pumps, valves, detectors, and power supplies [18]. This ancillary equipment makes these devices far from the portable, small, and low-power “labon-a-chip” ideal. Nevertheless, progress towards completely integrated, or at least partially integrated and highly portable, devices has been made. For example, portable microfluidic platforms used as biosentinels for space exploration have been described [19,20]. Handheld devices capable of rapid analysis of low concentration biowarfare agents have also been developed [21–23]. Furthermore, commercial manufacturers have made great strides in development of key components such as low current power supplies, lasers, light emitting diodes (LEDs) and pumps that are more amenable to portable microfluidic platforms. Integration can take the form of components that are microfabricated directly into the device, e.g. a microelectrode for electrochemical detection (ECD) that is patterned directly into a fluidic channel. Another form of integration is to use miniaturized components, fabricated separately, that are designed to “plug-in” as a module into the microfluidic device. An example is a photodiode that is embedded into a PDMS chip. In this approach, mating the components to the microfluidic manifold may require additional fabrication steps such as bonding (e.g. thermal or ultrasonic), gluing (e.g. UV activated epoxy) or packaging (e.g. flip-chip technology) [24]. In this review, we describe the trends in integrating key components such as detectors, pumps, and power supplies onto chips. We also describe some of the applications that are emerging for such devices. 2. Integration of detection transducers onto microfluidic devices Since their inception for chemical separations in 1993, microfluidic devices have been vigorously explored for their ability to separate a wide range of analytes using electrophoresis and chromatography resulting in a suite of techniques with impressive performance [25–28]. Separations performed on microfluidic devices can be performed extremely rapidly with migration times that range from 1 ms to 1 min [3]. Extremely high theoretical plate counts can be demonstrated on chips as well [29]. The typical injection volumes for a CE separation generally range from 1 to 50 nL, which enables researchers to analyze small volumes. It is also possible to integrate a number of discrete fluidic functions on a microfluidic device that yields advanced analytical functionalities. For example, on-line chemical derivatization using NDA and OPA followed by mixing and injection techniques have been integrated on microfluidic devices [30–32]. Furthermore, two-dimensional separations can be performed on microfluidic devices to generate greater peak capacities and higher resolution [7,29,33]. These developments

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demonstrate the potential for a variety of separations-based assays to be developed on chips. When performing separation analysis on chips, it is routine for the injection valves and separation channel to be integrated into a monolithic device. In contrast, it remains far from routine to integrate the detector onto the chip. This can be illustrated in Table 1 which summarizes the biochemical analyses performed on chips in the past 5 years. The table demonstrates applications to many classes of biomolecules including amino acids, biogenic amines, nucleotides, oligonucleotides, peptides, and proteins. A variety of detectors with varying detection limits have been demonstrated as well, as summarized by the table. While successful assays have been developed, the number of chips that utilize integrated detectors is small. For example, only 30% of the applications shown in Table 1 utilize an integrated transducing element. Of this 30%, the bulk of the integrated transducers are electrodes employed for ECD which are currently limited due to their poor limits of detection for biological analytes. True lab-on-a-chip functionality will require more integrated detection. Integration of detectors may also be expected to provide several other benefits. For example, within a microfluidic device it can be difficult to align detection elements with the appropriate fluidic stream; however, integration greatly simplifies this aspect of the instrument operation. Integration is also likely to facilitate detection in parallel systems where it is required to detect at multiple points on the chip. The degree and methods of integrating detector components is largely dependent on the fabrication method used. In the following section we will discuss common fabrication techniques and also how these methods have been utilized to integrate a number of elements pertinent to a wide variety of detection modalities including LIF, ECD, SPR, NMR and MS. 2.1. Fabrication techniques for transducer integration The three most popular fabrication techniques for microfluidic manifold generation are soft lithography [34], embossing/injection or casting molding [35], and complementary metal-oxide-semiconductor (CMOS) methods (e.g. micromachining using isotropic wet etching, reactive ion etching (RIE) and various chemical vapor deposition techniques) [36]. These techniques were used to create 90% of the total number of microfluidic devices reported over the past 14 years. Some recently developed techniques involving laser ablation and multiphoton instrumentation are beginning to be used for the development of monolithic three dimensional features which could provide new methods of integration that would obviate the aforementioned coupling techniques. Soft lithography using poly(dimethylsiloxane) (PDMS) is one of the most popular methods for microfluidic fabrication. The popularity of PDMS as a substrate for microfluidic fabrication largely originates from its surface energy, low cost, robust processing parameters, gas permeability, biocompatibility, and elastomeric siloxane backbone [16,37]. The high surface energy of PDMS enables it to bond to a wide variety of different surfaces via conformal contact between PDMS and the surface. For instance, PDMS can form a reversible conformal

bond with glass, metal, or photoresist that is strong enough to confine fluid within a microfluidic manifold without leaking. If pressure driven flows are applied to the microfluidic manifold, plasma or UV oxidation of the surface can be used to bond PDMS irreversibly to a variety of substrates. This sealing method maintains a leak free seal for moderately high pressures [38,39]. This characteristic can be used to place a variety of transducers, fabricated from non-PDMS materials, in conformal contact with a PDMS microfluidic manifold. The elastomeric nature of PDMS also enables PDMS to deform around an integrated transducer that may protrude several microns from the planar surface of the substrate, yet still retain a seal around the integrated component. These physical characteristics of PDMS reduce the number of fabrication steps required to build an integrated transducer capable of microfluidic analysis. Fig. 1a shows the integration of electrical components that were initially patterned on a glass slide and consequently bonded to a PDMS microfluidic manifold. The figure illustrates that the PDMS microfluidic channel can deform around an electrode. Embossing and injection molding technologies have also been used extensively to integrate a variety of optical components and transducers to microfluidic manifolds [40]. These fabrication techniques can mold thermoplastics around embedded features including emitter tips and fiber optics, illustrated in Fig. 1b. For embossing techniques, masters composed of silicon or metal are used as a template to mold micromachined features by pressing the template or heating it against a thermoplastic. Once the thermoplastic reaches its melting temperature (Tg ) it will form the shape of the template. Some popular substrates for embossing and injection molding include the thermoplastics: poly(methylmethacrylate), poly(carbonate), cyclic olefin copolymer, polystyrene, polyvinyl chloride, and polyethyleneteraphthalate glycol. This method is capable of forming microfluidic channels and micromachined guiding structures for the insertion of external transducer elements. Flow injection is similar to embossing in that it uses a thermoplastic, with the exception that the substrate is melted first and transported by injection into a chamber containing the template. The substrate is allowed to cool below the Tg yielding a matching template. Higher throughputs are common with injection molding over embossed substrates making this technique commercially attractive. Casting is also a useful method of microfluidic fabrication and sensor integration. Casting, in comparison to embossing and injection molding, provides a route to fabricate microfluidic manifolds with low cost and no mechanical force. This method has been used to integrate mechanical actuators, electrodes for heating, and electrodes for electrophoresis [24]. CMOS processing has also been used to develop microfluidic manifolds with integrated transducers in both glass and Si substrates. Silicon-based CMOS processing offers the possibility of integrating multi-level microfluidic channels with a variety of sophisticated electronics and optical features. Fig. 1c demonstrates the integration of an LED and a microfluidic channel on a monolithic device using CMOS processing [41]. Mastrangelo and co-workers [42] fabricated a microfluidic device with an integrated CMOS circuit that contained control, detec-

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173

Table 1 Biochemical species studied using microfluidic devices Chemical species

Detection mode

Integration

Substrate

LOD

Ref.

2-(Diethylamino) ethanethiol 2-(Dimethylamino) ethanethiol 6-Carboxy-fluorescein Actin Alanine

ECD

I

Glass

5 ␮M

[169]

ECD

I

Glass

5 ␮M

[169]

ICCD LIF CCD, LIF, ECD

NI NI I, NI, P–I

30 nM 30 nM* 8 nM–7.1 ␮M

[170] [33] [7,50,171,172]

LIF PAD Scanning-LIF ICCD ICCD PAD, LIF, scanning-LIF BEX-ICCD LIF ICCD CL CL Vis-absorption LIF

NI I P–NI P–NI P–NI I, NI, P–NI

Polycarbonate–PMMA PMMA PMMA, glass, PDMS, polycarbonate Glass, fused silica PDMS Glass PDMS PDMS PDMS, glass

100 fM–0.1 mg/mL* 5 ␮M 4.3 nM 2.0 mg/mL* 1.0 mg/mL* 10 pm–150 ␮M

[173–175] [176] [177] [178] [178] [7,30,172,176,179]

2 nM 60 nM–200 nM 1 mM* 1.5 ␮M* 3.8 nM 5–25 mg/mL 50 ng

[180] [6,7,11,30] [170] [181,182] [183] [184] [185]

NI NI NI, P–NI, I

0.1 mg/mL* 7.5–30 ␮M 100 fM–0.1 mg/mL*

[174] [186] [33,50,107,175,185–190]

Carbohydrate antigen Carbonic anhydrase Carcinoembryonic antigen Cardiac troponin 1

LIF Vis-absorption LIF, MALDI-MS, ECD, SPR ECD CCD, LIF ECD UV-vis, LIF

2 ng/mL 10 nM– 0.1 mg/mL* 2 ng/mL 30 ng/mL

[191] [173,174,192] [191] [193,194]

Catechol Cellobiose Chicken ovalbumin Cholesterol Chymotrypsinogen Chymotrypsinogen A Concanavalin A C-reactive protein Creatinine Crystal violet Cy5 Cytochrome C

ECD LIF LIF UV-vis, LIF Deep UV-LIF ECD LIF Light scattering, LIF ECD Vis-absorption LIF TLM, LIF, ECD

P–I NI NI NI NI P-I NI NI, I I NI NI NI, P–I

4.1 ␮M 1 ␮g/mL* 300 pg 7 ␮g/mL–25 mg/mL* 0.5 ␮M 30 ␮M* 30 nM* 0.1 ␮g/mL–1 ng/mL 150 ␮M* 0.54 ␮M 3.3 nM 1–30 ␮M*

[49] [195] [196] [197,198] [199] [50] [33] [194,200,201] [202] [203] [204] [50,189,205]

Deoxyuridine-5’triphosphate DNA

FCS

NI

PMMA PDMS, glass Polycarbonate–PMMA PDMS Glass Acrylic PDMS, poly(IBA), poly(HEMA), agarose, polycarbonate Glass, fused silica PDMS PDMS, glass, fused silica, PMMA–CD, polycarbonate Glass- Silicon PMMA, glass, fused silica Glass–silicon Polyacrylimide, PDMS, silicon Glass PMMA Glass PMMA Quartz Polycarbonate PMMA PMMA, silicon Glass Teflon PTFE Glass (Borofloat) Glass, PDMS, polycarbonate Glass

SMD

[206]

ICCD, DICM, LIF, Scanning-LIF ECD, AOD SVD, ECD, LIF FCS SVD LIF FTIR CCD

P–NI, NI, I

PMMA, PDMS, glass polycarbonate–PMMA

SMD—30 mM*

[17,35,162,165,170,192,207–235]

NI, I NI I NI NI NI

PDMS, quartz, glass PMMA PDMS–quartz Glass, fused silica Silicon PMMA

60 nM–125 ␮M 5 nM 500 nM, 1.2 ␮M 0.1 mg/mL* 0.115 M* 48 ␮g/L

[6,11,236–238] [239] [236] [174] [240] [236]

ECD FRET ICCD ICCD

I NI I I

Glass–silicon PDMS PMMA PMMA

2 ng/mL 0.48 nM 1 ␮g/L 0.7 ␮g/L

[191] [241] [242] [242]

Alcohol dehydrogenase Ampicillin Anti-estra-dial Anti-gp120 Anti-gp41 Arginine Asialoglycans Aspartate Aspartic acid ATP Atropine Blue dextran Botulinum toxin

Bovine erythrocyte Bromophenol blue BSA

Dopamine DsRed Epinephrine Equine liver Ethyl acetate Ethyl methylphosphonic acid Ferritin FLAG peptides Flavin adenine dinucleotide Flavin mononucleotide

P–NI NI NI NI I NI NI

I NI, I I NI

174

G.T. Roman, R.T. Kennedy / J. Chromatogr. A 1168 (2007) 170–188

Table 1 (Continued ) Chemical species

Detection mode

Integration

Substrate

LOD

Ref.

Fluorescein

LIF, UV-abs., ICCD, AOD-LIF

NI, P–NI

30 pM–10 ␮M

[7,170,177,232,243–246]

GABA GFP Glucose Glutamate Glutamic acid Glutamine Glycine

LIF AOD-LIF PAD, ECD LIF Scanning-LIF CCD, ECD LIF, ICCD, scanning-LIF

NI P–NI I NI P–NI I, P–I NI, P–NI

60 nM 1 ␮M* 90–100 ␮M 10–200 nM 3.3 mM* 10 nM–10 ␮M* 60 nM–1 mM*

[6,11] [246] [176,247,248] [11,30] [179] [50,192] [11,171,179]

Green fluorescent protein Helix pomatia lectin Histidine IgG IgG (mouse) IgG (rabbit) Immunosuppresive acidic prot. Insulin

AOD-LIF LIF PAD GMR sensors, LIF LIF LIF CL

NI NI I, NI I NI NI NI

PDMS, polycarbonate–PMMA, glass PDMS Glass PDMS, glass PDMS, glass Glass PMMA, polycarbonate PDMS, polycarbonate–PMMA, glass Glass (Borofloat) PMMA PDMS BCB, Glass PDMS PDMS Quartz

1 ␮M* 30 nM* 350 ␮M 7.0 pM 15.6 ng/mL 244 pg/mL 100 nM

[246] [33] [172,176] [249] [250] [250] [251]

LIF, ECD, ICCD, scanning-LIF TLS LIF CCD

NI, I, P–NI

10 pM–30 nM

[13,31,170,248,252]

NI NI I

Glass, polycarbonate–PMMA Glass Silicon PMMA

1 ng/mL, 60 pM 1 ng/mL 48 ␮g/L

[253] [200] [23]

SVD LIF ECD CCD SVD LIF ECD CCD, Deep UV-LIF, LIF LIF LIF PAD ECD CCD LIF CCD, LIF CCD, Deep UV-LIF, LIF, scanning-LIF LIF ECD ESI–MS, ESI–MS–MS LIF PAD LIF LIF CL ICCD, scanning-LIF

I NI I I I NI P–I NI

PDMS–quartz Glass Glass Glass–silica PDMS–quartz PMMA Polycarbonate PMMA, quartz, PDMS

515 nM 20 nM 100 ␮M* 1–2 mM 3.6 uM 30 nM* 30 ␮M* 10 nM–0.9 ␮M

[236] [173] [247] [254] [236] [33] [50] [188,192,199]

NI NI I P–I NI NI NI NI, P–NI

PMMA PMMA PDMS Polycarbonate PMMA PDMS, silicon PMMA, glass PMMA, quartz, glass

1 ␮g/mL* 1 ␮g/mL* 150 ␮M 30 ␮M* 48 ␮g/L 30 ng/mL 100 fM–10 nM 6.4 nM–30 nM*

[195] [195] [176] [50] [23] [194] [175,192] [33,171,177,255]

NI P–I NI

Glass Polycarbonate Cyclic olefin

300 pg 30 ␮M* 4 ␮g

[196] [50] [256]

NI I NI NI NI NI, P–NI

1 ␮g/mL* 5 ␮M 50 ␮M* 50 ␮M* 77 nM 500 ␮M*–1 mM*

[195] [176] [257] [257] [183] [170,179]

15 fmol 10 nM 1000 proteins/cell 48 ␮g/L

[190] [192] [259] [23]

56 ppm Single sphere 1.4 ␮M* 30 nM* 40 ␮M*

[258] [74,201] [255] [33] [260]

Interferon-␥ Interleukin-6 Isopropyl methylphosphonic acid Isoproterenol Lactalbumin Lactate Lactic acid l-Dopa Lectin peanut agglutinin Leu-enkephalin Lysozyme Maltose Maltriose Mannose Met-enkephalin Methylphosphonic acid Myoglobin Myosin Ovalbumin Ovomucoid Oxytocin P450 Panose Penicillin Peptide F-Calc Peptide F-Src Pethidine Phenylalanine Phosphoproteins Phosphorylase B Phycobiliprotein Pinacolyl methylphophonic acid Poly(ethylene glycol) Polystyrene ␮-spheres Propranolol Protein A Rhodamine B

MALDI-MS CCD BEX-ICCD CCD

P–NI NI P–NI I

PMMA PDMS PDMS PDMS Glass Polycarbonate–PMMA, glass PMMA–CD PMMA PDMS PMMA

RIG LIF, light scattering Deep UV-LIF LIF LIF

NI NI NI NI P-NI

PDMS Silicon, PMMA Quartz PMMA Glass

G.T. Roman, R.T. Kennedy / J. Chromatogr. A 1168 (2007) 170–188

175

Table 1 (Continued ) Chemical species

Detection mode

Integration

Substrate

LOD

Ref.

RNA R-phycoerythrin Serine

Scanning-LIF, LIF AOD-LIF LIF, ICCD

P–NI, NI NI NI

1 aM–50 pM 1 ␮M* 10 pM–60 nM

[162,216,261–263] [246] [6,7,11,170,264]

Serotonin Serum albumin Soybean trypsin inhibitor Staphylococcal enterotoxin B Streptavidin Sucrose Tacrolimus Taurine Tetanus neurotoxin Transferrin Trifluoperazine Trypsin inhibitor Trypsinogen Tryptophan

Deep UV-LIF CCD LIF LIF

NI I NI NI

PDMS–glass, silicon Glass (Borofloat) PDMS, glass, polycarbonate–PMMA Quartz PMMA Glass, fused silica Glass

1.9 ␮M* 10 nM 0.1 mg/mL* 28.5 fg/mL

[255] [192] [174] [265]

LIF SPR LIF LIF LIF LIF LIF CCD-LIF, LIF Deep UV-LIF CCD, Deep UV-LIF, ECD ICCD LIF

NI I NI NI NI NI P–NI NI NI NI, P–I

30 nM* 1% wt 5 ng/mL* 60 nM 2 nM 30 nM* 0.7 mM 100 fM–20 nM 0.5 uM 2.0–10 ␮M*

[33] [266] [267] [6,11] [268] [33] [269] [173,174,192] [255] [50,192,255]

NI NI

PMMA Glass PDMS PDMS Glass PMMA PMMA-CD PMMA, glass Quartz PMMA, quartz, polycarbonate Polycarbonate–PMMA PDMS, silicon

1 mM* 20 pg/mL, 1.14 pM

[170] [270]

LIF ECD LIF

NI I I

Glass Glass PDMS

300 pg 100 ␮M* 10 mg/L

[196] [202] [271]

CCD, LIF, ECD

NI, P–I

10 nM-10 ␮M*

[7,50,192]

LIF TLS ECD ICCD-FRET LIF CCD, Vis-abs., LIF, LIF ECD

NI NI I P–NI NI I, NI

30 nM* 40 nM 2 ng/mL 34 molecules/450 pL 100 ␮M* 100 fM–5 mg/mL

[33] [272] [191] [261] [32] [174,184,192,273]

I

PMMA, glass, polycarbonate PMMA Quartz Glass–silicon PDMS Glass PMMA, acrylic, mylar, glass Glass–silicon

2 ng/mL

[191]

AOD-LIF, MALDI UV-vis SPR

NI, P–NI I I

Glass, PDMS Silicon PDMS

1 ␮M*–2 mg/mL* 0.5 dl/mg 11 pM

[246,274] [88] [107]

Tryptophan Tumor necrosis factor (TNF␣) Turkey ovalbumin Uric acid Urinary human serum Albumin Valine Wheat germ agglutin Xylenecyanol ␣-Fetoprotein ␤-Actin cDNA ␤-d-Galactopyranoside ␤-Galactosidase ␤-Human choriogonadotropin ␤-Phycoerythrin Uric acid Glucose oxidase

If the transducer was located on the monolithic plane of the microfluidic device it was classified as integrated (I) and if it was located off the plane of the microfluidic device it was classified as non-integrated (NI). Transducers that were configured for parallel analysis are either P–NI or P–I, for parallel non-integrated and parallel integrated, respectively. Abbreviations: PS: polystyrene, TLS: thermal lens spectroscopy, DICM: differential interference contrast microscopy, SMD: single molecule detection, AOD-LIF: Acousto-optical deflection laser induced fluorescence, CL: chemiluminescence, ISE: ion selective electrode, FCS: fluorescence correlation spectroscopy, RIG: refractive index gradient, ECL: electrogenerated chemiluminescence, PAD: pulsed amperometric Detection, SVD: sinusoidal voltammetric detection, CCD: indirect contact conductivity detection, ICCD: integrated contact conductivity detection, PMMA: poly(methylmethacrylate), PDMS: Poly(dimethylsiloxane), GMR: giant magnetoresistive, BCB: bisbenzyocyclobutadiene, IBA: isobornyl(acrylate), HEMA: 2-hydroxymethyl methacrylate.

tion, drive, and communication electronics. The fabrication was performed by using a CMOS device that contained all the aforementioned integrated components that was further processed by adding a series of layers that were sacrificially etched to produce a microfluidic manifold. Unlike the previous fabrication techniques, sacrificial etching was used to generate microfluidic channels which requires no thermal or pressure bonding procedure. This method facilitates the micromachined alignment of both optical transducers and fluidic channels for example. The avoidance of thermal and pressure driven bonding steps assists in coupling electronics that would be damaged with the extreme pressures or temperatures. Furthermore, transparent sil-

ica layers, or polymers can be used rather than Si which provide microscopic observation. Peeni et al. [43] have recently reviewed a number of sacrificial layer fabrication methods using CMOS for microfluidic devices. These new methods have the potential of fabricating both low cost and high quality integrated microfluidic devices that couple both electrical interfaces and fluidic systems on glass or silicon substrates. Finally, three-dimensional lithographic techniques are beginning to be explored which have the capability of integrating a variety of optical components. Schmidt et al. [44] has demonstrated a two-photon 3D lithographic technique for the fabrication of waveguides over a printed circuit board that con-

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Fig. 1. Integration techniques on microfluidic devices: (A) soft polymer lithography used for integration of electrodes, (B) PMMA embossing used for the coupling of fiber optics and emitter tips, (C) CMOS processing for LED and photodiode integration, and (D) two-photon three-dimensional lithography used for waveguide generation and registration to a variety of features including LED, photodiode and microfluidic channels.

tains a laser and photodiode. This method allows high precision registration of waveguides with lasers and photodiodes. As with sacrificial layer fabrication methods this technique also reduces the bonding steps needed in a fabrication process and hence could improve throughput for inexpensive mass production of integrated microfluidic devices. This technique holds great promise for further development of totally integrated microfluidic devices (Fig. 1d). 2.2. ECD integration The most popular integrated transducer for microfluidic devices are those used for ECD, as can be seen in Table 1. This popularity stems not only from the fact that the transducer must be within close proximity of the microfluidic channel, but also due to the ease of fabrication processes required for electrode fabrication. ECD requires the integration of electrodes on-chip that is either in direct or indirect contact with the microfluidic channel. Electrodes can be used for a variety of different ECD modes including amperometry, conductivity, and potentiometry. A number of reviews have recently discussed the integration of these electrochemical techniques on microfluidic devices [45–47]. Although a number of electrochemical methods have been developed for single-point detection, only a few examples of parallel and/or portable analysis using ECD have been demonstrated. Ferrigno et al. [48] developed a 10-electrode potentiometric array on a PDMS device for the calculation of redox potentials for electrochemically active ions. Mathies and co-workers [49] developed multiple electrodes on a microfluidic device for catechol quantitative measurements. More recently, Shadpour et al. [50] have fabricated a parallel device capable of contact conductivity detection on a polycarbonate microchip. Several other researchers have employed multiple electrodes on a single channel for electrochemical decoupling, but this technique has yet to be multiplexed. As

illustrated in Table 1, a disadvantage of ECD includes relatively poor sensitivity. Electrochemical methods on microfluidic devices typically have detection limits worse than 100 nM. This poses problems for experiments that aim to detect trace level components in biological systems. The modest concentration LODs are somewhat surprising given that in HPLC and capillary LC systems, the detection limits achieved by amperometric detection can be 1 nM or better. More research is required to determine the reasons for the shortfall by microfabricated systems. It seems unlikely to be an inherent limitation of ECD. 2.3. Fluorescence detection integration Table 1 illustrates that LIF is by far the most popular detection method for biochemical analysis in microfluidic systems. This popularity derives from its high sensitivity for low concentration analytes and its ability to be coupled with separation modalities that can analyze a variety of compounds in complex mixtures. Despite the excellent sensitivity and utility of LIF detection, the integration of LIF detection into microfluidic devices has lagged tremendously behind the progress of ECD integrated microfluidic chips. The reason for the limited development of LIF integration is due primarily to the large number of optical, electrical, and transducer components that must be integrated to generate a complete system. The components for LIF detection include light sources, light collection and delivery optics, filters and photon detectors. A long-term goal for the field may be to fully integrate these components into chips for a fully functional and integrated detection system. An intermediate goal is the development of a system in which some or all of the detector components are external to the fluidic chip, but are highly miniaturized and designed to allow easy alignment. Such “plug-in” components would be useful with disposable fluidic chips and could be highly portable.

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2.3.1. Integrated excitation sources Commercially available LEDs with high power and narrow bandwidths spanning the near-UV to IR spectral range are beginning to be implemented for integrated microfluidic fluorescence detectors. These relatively small (sub-millimeter) and inexpensive light sources can be combined with PDMS soft lithography to fabricate disposable integrated devices using a plug-in approach. Although these external optical components can be successfully bonded, glued or imbedded next to a microfluidic channel, they are often much larger than the microfluidic device. Further miniaturization techniques aim at developing fully packaged and functional integrated light sources that do not require plug-in of external optical components. In addition to improving miniaturization, this process has the potential of further improving limits of detection by reducing the layers that the light must travel through and hence the number of reflections, or the amount of light scatter that is generated [51]. Only a handful of fully integrated light sources have been developed for monolithic microfluidic devices. These light sources include: (1) LEDs [52–56], (2) carbon nanotube FETs [57], (3) lightemitting nanofibers [58] and microdischarge [59]. Choudhury et al. have developed an organic light-emitting device (OLED) on a microfluidic device for the detection of glucose. The OLED was fabricated by vacuum evaporation of organic layers on an indium tin oxide (ITO) coated glass substrate [55]. Edel et al. have fabricated polymer light emitting diodes for microscale capillary electrophoresis. The LEDs were fabricated using an ITO coated glass substrate followed by a coating of poly(3,4ethylenedioxythiophene)/polystyrene sulfonate [PEDOT/PSS] [56]. Fig. 2 illustrates a liquid electrode discharge spectral emission chip (LEdSpEC). This device integrates a microfluidic wavelength-tunable optical source with a microfluidic manifold for biochemical detection. A discharge is produced in air between a cathode and anode that is filled with a metal ion solution. A line spectra is generated from the ionization of the metal ions and is consequently filtered and guided to the biochemical sample. A barium chloride solution was employed to generate wavelengths of 454 and 493 nm which was suitable for DNA fluorescence. Lead (II) nitrate was employed to provide a 280 nm

Fig. 2. Illustrates a microfluidic discharge-based optical source that is fabricated on a stacked microchip. The abbreviated arrows illustrate the stacking whereas the longer arrows indicate the optical path of the emission light that is emitted from a biochemical sample. [59] – Reproduced by permission of The Royal Society of Chemistry.

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emission line which could be used for tryptophan fluorescence [59]. This method is useful in that it provides a miniaturized method of producing deep UV light for biochemical analysis of amino acids such as tryptophan, tyrosine or phenylalanine. The bulk of the light sources mentioned previously are limited to near-UV and visible light generation. Optically pumped dye doped liquid waveguide lasers have also been developed to provide a number of integrated light sources on-chip [60–63]. Using solutions of rhodamine 640 and methanol it was possible to generate wavelengths between 615 and 630 nm by changing the methanol concentration. The maximum output pulse energy of these waveguides was 25 ␮J. These devices require pumping sources which are often LEDs. To date, dye-doped liquid core waveguides have yet to be implemented in biochemical investigations. LEDs integrated on microfluidic devices have been demonstrated in proof-of-principle separations of purified DNA fragments at 100 ng/␮L, [64,65] but have yet to be demonstrated for quantitative analysis of biological samples. 2.3.2. Integrated waveguides and fiber optics Non-integrated light sources have recently been coupled to microfluidic chips using integrated fiber optics and waveguides [40,66–69]. Fiber optics have been integrated on-chip using both soft lithography [68] and embossing [40] procedures. The elastomeric nature of PDMS, as described above, enables researchers to plasma bond and cure microfluidic manifolds in close proximity to fiber optics [35,68]. PMMA embossing has also been used to integrate fiber optics with microfluidic manifolds. This is performed by initially embossing the microfluidic channel manifold and the fiber optic channels simultaneously. Following embossing the fiber optic is manually inserted into the channel and aligned. The cylindrical geometry of the fiber optic does not mate with the square geometry of the embossed channel. To rectify this problem, a sealing process is performed by applying moderate pressures and temperatures to both the manifold and a blank PMMA cover. The process slightly melts both the fluidic manifold and the blank causing the normally square channels to round. The same phenomenon is present in rounding of positive photoresists for the fabrication of actuation valves [70]. Waveguides have also been integrated on-chip for delivery and collection optics [69,71–73] and are beginning to be used for multiple point detection on a microfluidic manifold [74,75]. Solid or liquid core waveguides have been developed using a number of fabrication techniques and materials including glass [75,76], SU-8 [69,77], PDMS [71] and mesoporous materials [78]. Jiang and Pau describe a method of integrating waveguides using a microfluidic channel that can be configured in a spiral geometry. Fig. 3 illustrates a SU-8 core multimode spiral waveguide that is 40 ␮m thick and 50 ␮m wide. This waveguide is used for spectroscopic analysis and can act as a compact optical sensor [69]. Mogensen et al. have also developed a number of SU-8 core waveguides that have been integrated on microfluidic manifolds. These waveguides are being multiplexed to fabricate a device capable of calculating, via Fourier transform, the velocity of particles within a microfluidic manifold [73,74]. There are

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Fig. 4. Illustrates an integrated long-pass filter, microfluidic manifold, an organic photodiode, and an OLED. Sudan II was added to the PDMS prepolymer prior to polymerization yielding a dye doped PDMS manifold shown to the right. A second non-doped PDMS flat slab was used to seal the microfluidic manifold. Following sealing, co-linear optical geometry was established for the delivery of excitation light using an OLED and detection using an organic photodiode. [89] – Reproduced by permission of The Royal Society of Chemistry. Fig. 3. Illustrates on-chip spiral waveguides that were fabricated using a microfluidic manifold. Since these waveguides are fabricated from polymers they are disposable and potentially low cost. [69] – Reproduced by permission from the American Institute of Physics.

several limitations of solid and liquid core waveguides. Some of these include light scattering, efficiency loss over extended lengths and non-UV transparency or high autofluorescence for the core materials. Because fiber optics and waveguides enable micromachined optical alignment, these devices can be used in disposable applications that do not require time-consuming alignment procedures. In addition, miniaturization of the detection optics over conventional LIF objectives is also advantageous as it reduces the overall size of the device. 2.3.3. Integrated lenses and filters Many of the light sources integrated onto monolithic devices are non-collimated and must be focused for efficient fluorescence detection. Ball lenses can be plugged into the chip to help focus the light; however, it is impossible to focus this light to a Rayleigh limited waist diameter ∼λ/2 which can severely impact the limits of detection as compared to confocal setups using coherent laser lines and external objectives [68]. In attempts to further focus non-collimated light sources a variety of lenses have been fabricated directly within the chip [79–83]. These lenses have provided some improvement in detection efficiency and fidelity by producing 50 ␮m waist diameters. Focusing also has the potential of generating a more symmetrical excitation volume across the width of the channel as compared to illumination from butt ended waveguides or fiber optics that widen across a channel. Future improvements in on-chip lasers for coherent light generation and the optics required to focus this light to a Rayleigh limited waist diameter are needed if integrated fluorescence detection is going to be successfully developed for highly sensitive biochemical systems. Such advancements would provide an integrated means of LIF detection that is comparable or equal to conventional non-integrated optics and lasers. On-chip filters have been developed to assist in separating excitation from emission wavelengths. These filters are either integrated monolithically to the transducer [65,84–88] or the substrate [89,10,90–92]. Filters that have been integrated

monolithically include Fabry–Perot filters or absorptive layers deposited directly over the transducer. For instance, Burns et al. [65] have deposited 40 alternating SiO2 /TiO2 layers over a microfabricated photodiode to generate a long-pass filter at ∼500 nm. Applications of transducer coating methods are largely limited to non-disposable devices as they are expensive to fabricate. Other methods involve coating or doping the microfluidic substrate with molecules that block short wavelength light, and yield long-pass filters. These methods are amenable to soft polymer lithography and are useful in applications that necessitate disposability. For instance, Hofmann et al. have added lysochrome dyes to a PDMS prepolymer prior to molding, illustrated in Fig. 4. The dyes were mixed and homogenously distributed in the prepolymer followed by molding to an SU-8 mold. After removal of the polymer from the mold, it was consequently analyzed optically and demonstrated to have 0.01% transmittance below 500 nm and >80% transmittance above 570 [89]. The simplicity of fabrication for these dye doped PDMS filters is extremely attractive for the development of disposable devices that have the potential to be used in point-of-care applications. 2.3.4. Integrated light detectors A variety of transducers have been integrated on microfluidic devices for LIF detection including: photodiodes [41,64,93,54,94–96], a CMOS imager [97], and a crystalline silicon detector array [98]. Kamei et al. have developed an amorphous-Si–H pin photodiode using PECVD which is illustrated in Fig. 5. Due to the low processing temperature of ∼200 ◦ C this device can easily be integrated on low cost glass or plastic substrates [64,93]. This fully integrated photodiode was used to detect fluorescein at 17 nM and a variety of DNA fragments. Alternatively, photodiodes fabricated using c-Si must be fabricated at 1000 ◦ C and hence expensive quartz substrates must be used [98]. In addition, a-Si–H films have absorptive coefficients that are an order of magnitude greater than c-Si films in the spectral region between 500 and 650 nm. The dark currents of a-Si–H films are on the order of 10−11 A/cm2 which also provide better limits of detection than c-Si [99,100]. Wang and Hofmann et al. [41,101] have integrated organic photodiodes using cop-

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Fig. 5. A schematic of a hybrid annular shaped Si–H photodiode for the detection of fluorescence emission from a microfluidic channel. A ZnS/YF3 optical longpass filter was deposited on top of the photodiode to filter the excitation from the emission light. An argon ion laser was focused onto a microfluidic channel using a ball lens. [93] – Reproduced by permission from the American Chemical Society.

per phthalocyanine-fullerenes and poly(3-hexylthiophene) with [6,6]-phenyl-C61-butyric acid-methylester. These photodiodes were integrated to detect chemiluminescence assays of a peroxyoxalate reaction on microfluidic devices. Microavalanche Si photodiodes have also been embedded in PDMS microfluidic devices [68], while miniaturized spectrometers employing CMOS imagers and crystalline silicon detector arrays have been developed [88]. 2.3.5. Other detectors Other techniques are beginning to emerge that aim at providing label free detection. One of the disadvantages of LIF is that analyte must be fluorescent to be detected. If the sample is not natively fluorescent, then a series of chemical processing steps must be performed to fluorescently derivatize the sample. A number of derivatization methods have been reported and reviewed; but the utility of derivatization can be limited by the time required for reaction or by the lack of reagents suitable for the functional groups of interest [102]. These chemical steps also occupy space on the microfluidic device and serve to increase its lateral dimensions. Therefore, label free detection is of significant interest. Surface plasmon resonance aims at eliminating the need to perform labeling procedures by providing a “universal” detector. Detection systems that employ surface plasmon resonance operate by monitoring the dielectric change near the region of a gold surface. This dielectric change can occur by either a bind-

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ing event to the surface of the gold via an antibody, or simply a dielectric change that occurs within 200 nm of the gold surface [103,104]. SPR sensor systems have been integrated on PDMS microfluidic devices by placing a PDMS manifold in direct contact with the SPR sensor. These integrated devices have been demonstrated to be capable of detecting cardiac markers, BSA, streptavidin, protein A and IgG with a concentration range from 5 pg/mL to 100 ng/mL [105,106]. The detection of these analytes has been performed solely via immunological reactions at the gold surface of the transducer. The first non-immunological detection of analytes on a microfluidic-SPR device has recently been demonstrated by Ly et al. This device consists of an 8 mm × 500 ␮m × 500 nm separation channel bonded to a series of gold patterned surfaces. Pinched injections and separations of BSA and GOx on an SPR coupled microfluidic device have been performed at concentrations ranging from 30 to 11 pM, respectively [107]. Field effect transistor (FET)-based biosensors, fabricated on IC circuits, are beginning to gain popularity for integrated microfluidic-based experiments due to their operational simplicity and CMOS fabrication protocols that are capable of mass production. Such sensors have been integrated onto microfluidics chips for detection of glucose [108], oxygen [109], and streptavidin–biotin [110]. Modified FET sensors called extended gate field effect transistors (EGFETs) are popular due to the isolation of the FET from the chemical environment using a membrane that is chemically sensitive. The advantages of such a device over a simple FET detector includes both heat and temperature insensitivity [111]. The results of integrating more information-rich detectors, or at least some of their components, are also starting to appear. Solenoid microcoils have been fabricated against glass microfluidic manifolds for use in NMR detection. Such integrated coils have been used to determine neat organic solutions, but have yet to be utilized for biochemical analysis [112]. Finally, a number of micro-ion-trap mass spectrometers are beginning to be fabricated suggesting the potential of developing a miniaturized mass spectrometer on a chip that can be coupled with microfluidics and CE [113]. Non-integrated coupling of MS has been widely attempted with varying success by integrating emitter tips using embossing and flow injection of PMMA and cyclo-olefins [114,115]. 3. Integration of pumps and power supplies for fluid manipulation In addition to detection methods, it is also necessary to consider integration of stable pumping and valving methods that enable not only fluid control on-chip, but also sampling from external environments. Conventional pumping methods require external equipment including syringe or peristaltic pumps and power supplies. These conventional methods are too bulky and consume too much power to be used in portable applications. Furthermore, at low flow rates syringe pumps have been demonstrated to generate significant oscillations due to vibrations on the walls of the pumping mechanism and a variety of effects due to non-linear friction [18]. As a result of these limitations, there

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have been tremendous efforts over the past decade to develop micropumps. There are essentially two types of micropumps: (1) reciprocating pumps and (2) continuous flow pumps. In reciprocating pumps there is an actuator, pump chamber and an inlet-outlet. The actuator is modulated using a variety of physical forces including electromagnetic, piezoelectric, thermopneumatic, pressure, and electrostatic forces. Continuous flow pumps have been demonstrated using a number of physical forces including: electroosmotic flow, ultrasonic waves and magnetohydrodynamic forces. Continuous flow pumps are generally easier to fabricate due to the lack of moving parts and can also generate non-pulsatile flows as compared to reciprocating pumps. A variety of non-conventional microfluidic pumping methods have recently been published that employ either continuous flow or reciprocating mechanisms. These include electrokinetic [116–125], microfabricated peristaltic [70], piezoelectric actuated [126–129], optically driven [130–132], ferro-fluidic magnetic [133–135], capillary [136], bubble [137,138], osmotic [122,123,139–142], electrohydrodynamic [143], electrostatic [144], thermopneumatic [145,146], nozzle-diffuser [147], conductive polymer [148,149], fluid responsive polymer (e.g. hydrogel) [150], electro-rheological [151], magnetohydrodynamic [152], buoyancy driven [153], and passive gravity-based [154] pumps. The integration of these methods on a fully monolithic microfluidic device will be important for advancing the development of microfluidic devices towards fully autonomous platforms. Some of these micropumping methods are better suited for portable application than others. We will further elaborate on robust micropumps that consume limited power and require limited auxiliary equipment as these methods are more likely to be employed in portable and completely integrated microfluidic devices. In addition, pumps that are compatible with biochemical analysis must also be capable of sampling solutions from external environments. Often this requires moderate backpressure capabilities. Microfluidic pumps that have shown promise for integrated operation on chips include electrokinetic pumps and microfabricated pumps actuated by either gas pressure or piezoelectric devices. These methods have been thoroughly characterized, require limited power consumption, and can pump against moderate to high backpressures. Many of the auxiliary components for these microfluidic pumps, notably electrokinetic and piezoelectric-actuated pumps, can be miniaturized to lateral dimensions similar to that of other components (e.g. a-Si–H photodiodes), making them extremely useful for microfluidic integration. Furthermore, electrokinetic pumps have been demonstrated to generate stable flows over time for capillaries with suitable surface chemistry and hence can be useful for separations and non-mechanical valving. 3.1. Electrokinetic pumps Electroosmotic flow has been known since the mid 70s [155]. More recently electroosmotic flow has been implemented on monolithic microfluidic devices for electrokinetic (EK) micropumps. EK micropumps were originally reported by Harrison and co-workers [5,125] who demonstrated the pumping of a

variety of analytes for chemical reactions and separations. Electric field control in such pumping mechanisms allows the facile control of both flow rate and polarity. EK pumping is also independent of the hydraulic diameter of the channel. This compares favorably with pressure driven flows were the flow rate is inversely proportional to the square of the hydraulic diameter [156]. Furthermore, EK pumps can also provide reproducible low flow rates (sub-␮L/min) at high backpressures as compared to mechanical micropumps, which may suffer from leakage from check valves and dynamic sealing of pistons [122]. EK pumps can be used with either a packed column or open capillaries. Packed capillary bed EK pumps developed by Chen et al. have been demonstrated to be capable of generating up to 2800 psi at flow rates of 1.6 ␮L/min [120,157]. Open channel EK pumps have been fabricated on microfluidic devices by Lazar and Karger [141] with flow rates that range from 10 to 400 nL/min with pressures up to 80 psi. Unfortunately, there are several limitations of EK pumping. The most significant is that changes in surface chemistry of the channel or composition of the fluids being pumped modify the zeta potential and hence the flow velocity. It also becomes problematic to incorporate mechanical valves in channels that have electric fields applied across the length of the channel. To combat these limitations, electric field-decoupled microfabricated EK pumps have been developed [120]. Overall, this method is advantageous over mechanically driven flow for several reasons including the improvement in the effective pressures and the general ease in implementation. 3.2. Microfluidic peristaltic membrane and piezoelectric pumps Two reciprocating pumps that are beginning to receive a great deal of attention are microfabricated peristaltic gas actuated pumps and piezoelectric micropumps. These two methods are capable of integrating a microfabricated actuation source on a microfluidic manifold that produces moderate pressures. Unger et al. have developed a method to perform peristaltic pumping by applying pneumatic pressure to a control channel located in one layer of a PDMS chip that causes deflection of a thin membrane and hence occlusion of a fluidic channel in a lower layer. Pressure application in sequence of control channels positioned along a fluidic channel results in peristalsis in the fluid channel [70]. These microfluidic pumps are capable of low flow rates as low as 10 nL/min and have been used for sampling from biological microenvironments [158]. Grover et al. [159] have developed pneumatic peristaltic pumps and valves that are capable of low dead-volume pumping from external probes. While similar in principle to the pumps described above, these pumps utilize primarily glass chips with only a thin PDMS layer to provide flexibility for the valve actuation. The latter system has advantages in minimizing sample and fluid contact with PDMS which can adsorb/absorb chemicals. These valves and pumps in combination with CE–LIF has been utilized for portable microfluidic analysis of amino acids [160]. Although these methods have proven useful for efficient pumping on-chip, the auxiliary components, including com-

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Fig. 6. Illustrates an integrated flow controller with piezo microvalve and flow meter printed on a circuit board. [275] – Reprinted by permission from Springer.

pressed gas and solenoid valves, are configured in a “plug-in” format rather than fully integrated on the chip. The plug-in format and the size and power consumption of the external solenoids and gas chambers dramatically increases the overall size of the device and its autonomous capabilities. Piezoelectric actuation sources are fully integrated on the microfluidic device and in combination with elastomeric polymers can be used to integrate self-contained peristaltic pumps in an integrated format shown in Fig. 6 [128,129]. Such pumps have been used for relatively high flow rates that range from 20 to 300 ␮L/min. 3.3. Electrical power supplies and control The aforementioned pumping methods all require electrical power supplies and electronic control. A number of power and electrical control methods have been developed for microfluidic manipulation and analytical detection instrumentation. Several advanced electrical control methods have been developed by Sandia National Labs. Fruetel and co-workers have developed a high-voltage module for electrokinetic separations. In addition, on-board, menu-based software enabled control of the instrument through an LCD screen and user interface. On-board microprocessors and memory also stored electropherograms and other operating conditions for later analysis [21,22]. These electrical components are contained within a module that is 7 in. × 8 in. × 4.5 in., weighing approximately 2.7 kg. Jackson et al. [167] have developed a miniaturized, battery-powered, high voltage power supply that also couples both ECD and interface circuits shown in Fig. 7. This device can operate continuously for 15 h without charging at 1 kV and 380 ␮A. The CE power supply is connected to the microchip through an interface circuit consisting of via relays, resistors and diodes. The data for this device is recorded using commercially available interface cards and a laptop computer. The combined size is 4 in. × 6 in. × 1 in. with a weight of 0.35 kg. The electronic and control systems described here can be used to quantitatively analyze biochemical systems using either ECD or LIF detection in a portable format. 4. Complete systems for biochemical analysis Over the past decade, several partially integrated microfluidic devices have been developed with the goal of analyzing

Fig. 7. Illustrates a miniaturized electronic control system for capillary electrophoresis and ECD. The integrated components include: (a) high voltage power supply, (b) an interface used to control the power supply and (c) the ECD detection circuitry. [167] – Reproduced by permission from the American Chemical Society.

biochemical samples including DNA, RNA, proteins, peptides, amino acids, neurotransmitters and a variety of metabolites in a portable environment. A majority of these partially integrated microfluidic devices are capable of plugging into a self-contained module that provides either fluidic or analytical control. Table 2 summarizes the salient features of these stateof-the-art devices that represent a high level of integration. In this section, we discuss some of these devices and their potential applications. 4.1. DNA and RNA Various forms of DNA and RNA analysis are extremely attractive targets for fully integrated systems. Such analyses have application in forensics, diagnostics, biowarfare agent detection, and gene sequencing. As a result, many examples of integrated systems directed towards such analyses have been described. Liu et al. have recently developed a portable PCR–CE microfluidic device for the genetic analysis of rapid short repeats for forensic analysis. This device integrates PCR thermal cycling, capillary electrophoresis separation, and elastomeric valves for pneumatic fluidic control. The off-chip components for this device are configured in a portable “plug-in” platform and include the LIF detection components and control/power systems. The device had an analysis time of 1.5 h and was capable of quantifying 20 DNA copies in a sample [161]. A similar device was configured in a parallel format for the analysis of biomarkers

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Table 2 Portable microfluidic devices for biochemical analysis Chemical species

Detection mode

Substrate

Chip footprint

Instrument footprint

LOD

Ref.

Alanine Aminoisobutric acid Arginine Aspartic acid BSA Carbonic anhydrase Catechol CCK peptide DNA Dopamine Glutamic acid Glutamic acid Glycine HPTS dye IgG IgG Ovalbumin Ovalbumin Proline Ricin Serine Serine Staphylococcal enterotoxins Valine

LIF LIF LIF LIF LIF LIF ECD LIF LIF ECD LIF LIF LIF LIF LIF LIF LIF LIF LIF LIF LIF LIF LIF LIF

Glass–PDMS Glass–PDMS Glass Glass–PDMS Glass Glass Glass Glass Glass–PDMS Glass Glass–PDMS Glass Glass–PDMS Glass Quartz Glass Quartz Glass Glass Quartz Glass–PDMS Glass Quartz Glass–PDMS

10 cm diameter 10 cm diameter 5.1 cm × 2.5 cm 10 cm diameter 2.0 cm × 2.0 cm 2.0 cm × 2.0 cm 5.1 cm × 5.1 cm 2.0 cm × 2.0 cm 2.0 cm × 7.5 cm 5.1 cm × 5.1 cm 10 cm diameter 5.1 cm × 2.5 cm 10 cm diameter 2.0 cm × 2.0 cm 5.1 cm × 7.6 cm 2.0 cm × 2.0 cm 5.1 cm × 7.6 cm 2.0 cm × 2.0 cm 5.1 cm × 2.5 cm 5.1 cm × 7.6 cm 10 cm diameter 5.1 cm × 2.5 cm 5.1 cm × 7.6 cm 10 cm diameter

36 cm × 36 cm 36 cm × 36 cm 28 cm × 16 cm 36 cm × 36 cm 12 cm × 12 cm 12 cm × 12 cm 7.6 cm × 10 cm 12 cm × 12 cm 21 cm × 31 cm 7.6 cm × 10 cm 36 cm × 36 cm 28 cm × 16 cm 36 cm × 36 cm 12 cm × 12 cm 13 × 7.6 cm 12 cm × 12 cm 13 cm × 7.6 cm 12 cm × 12 cm 28 cm × 16 cm 13 × 7.6 cm 36 cm × 36 cm 28 cm × 16 cm 13 cm × 7.6 cm 36 cm × 36 cm

0.4 ppb* 5 ppb* 2 ␮M 0.13 ppb* 1 nM 1 nM 4.6 mM* 1 nM 1.0 aM 2.2 mM* 0.07 ppb* 2 ␮M* 0.2 ppb* 10.0 pM 200 nM* 1 nM 200 nM* 1 nM 2 ␮M* 200 nM* 0.4 ppb* 2 ␮M* 200 nM* 130 pM

[160] [160] [20] [160] [21] [21] [275] [21] [65,160] [167] [160] [20] [160] [21] [22] [21] [22] [21] [20] [22] [160] [20] [22] [160]

and gene expression. This device was capable of analysis times of 45 min with attomole detection limits [162,163]. Blazej et al. fabricated a bioprocessor capable of integrating all of the components for Sanger DNA sequencing. The microfluidic bioprocessor and its integrated fluidic components are illustrated in Fig. 8. The integrated components include thermal cycling, sample purification and capillary electrophoresis separation. The off-chip components for this device include pneumatic control lines, LIF detection and electronic/power control systems [164]. An integrated photodiode and capillary electrophoresis system has also been developed by Kamei et al. [93] which was consequently used to detect DNA fragments and fluorescein with low nM detection limits. Overall, these portable devices could offer a facile method for point-of-care gene sequencing. Easley et al. [165] have also demonstrated an integrated genetic analysis system. This device has a number of integrated sample handling components including: nucleic acid purification using solid phase extraction, PCR using thermal cycling and elastomeric valves, and electrophoretic separation. The off-chip components for this device include the LIF detection components, syringe pumps for pressure driven fluid flow, IR thermal cycler, and control/power systems. This device was capable of performing total analysis of DNA in under 24 min with separation migration times of less than 180 s. The authors speculate that their device plausibly has a detection limit on the order of a few hundred starting copies of DNA. These devices have the capability of providing forensic analysis of DNA at a crime scene. The portability, rapid analysis, and low sample volume requirements would enable detectives to further accelerate crime investigations with improved accuracy.

Pal et al. have developed an integrated microfluidic device that is capable of genetically analyzing viral DNA (e.g. influenza). This microfluidic device integrates fluidic manipulation, thermal heaters, temperature sensors, valves and electrophoretic separations. Off-chip components include the pneumatic control, epi-confocal illumination, ICCD camera, and power/control systems. Separations of the PCR products were performed in under 140 s while 35 thermal cycles were performed in under 20 min. This device is fabricated from a silicon substrate, which could provide future advancements including the integration of LIF optics and CMOS control units [17]. Furthermore, this device has the capability of analyzing a number of pathogens including influenza. The integration and potential future portability of such a device could enable a variety of pointof-care analysis for influenza at airports, hospitals, or areas of mass transit. 4.2. Proteins and peptides Protein and peptide analysis on integrated chips is of interest for diagnostics, biowarfare agent detection, proteomics, and many routine biochemistry assays. Renzi and Fruetel et al. have developed portable microfluidic devices which can be used for the detection of biotoxins including ricin, staphylococcal enterotoxin variants, IgG, and ovalbumin. The microfluidic device integrates a number of sample handling features including a fill/flush port with filter, an electrokinetic pinched injector, and a separation channel used for capillary electrophoresis. The technological advancement of this device resides primarily in the hand held module that contains all of the LIF detection components, electrokinetic pumps and electronic control and power

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included the electronic and control systems which enable this device to be fully portable [50]. Vieillard et al. have developed an integrated microfluidic device with optical waveguides and capillary electrophoresis channels for the detection of a number of proteins including ␤-lactoglobulin A and carbonic anhydrase II [166]. 4.3. Amino acids, neurotransmitters, and small metabolites

Fig. 8. Illustrates the integrated fluidic components for Sanger DNA sequencing. Each monolithic device contained two complete bioprocessors capable of Sanger DNA sequencing. Shown in (A) is one half of the complete device. Shown in (B–F) are the individual microfluidic functions that are integrated on the device. [164] – Reproduced with permission from the authors and the Proceedings of the National Academies of Science.

systems. For operation, this device only requires the pipetting of labeled sample to analyze directly into the microfluidic device. Separations on this device were performed between 90 and 350 s. Electrokinetic micropumps were also powered and controlled using power supplies and control equipment integrated into the hand-held module [21,22]. The fluidic chip has a total footprint of 2 cm × 2 cm, whereas the handheld module has a footprint of 11.5 cm × 11.5 cm. This device has the potential of being further developed into a portable device for biotoxin detection which could be used in scenarios where portable and time critical elements are important. Further advancements must be made by coupling this device to a condenser capable of transferring biowarfare agents from a gaseous phase to a liquid phase for downstream analysis. Finally, both labeling and sample pumping must be performed on-line to generate a completely automated device. Shadpour et al. has developed a contactless conductivity detector integrated to a microfluidic device for the parallel detection of a series of proteins including leucine enkephalin, methionine enkephalin and oxytocin. A total of 16 channels were integrated with conductivity electrodes. Off-chip modules

Amino acids, neurotransmitters and metabolites are an attractive target for many microfluidic devices due to their importance in a variety of biological systems. Such devices have the potential of further characterizing metabolomic diseases, searching for extraterrestrial life, serving as biosentinels for space exploration, and investigating neuroscience phenomena just to name a few. Jackson et al. have developed a microfluidic device for the ECD of dopamine. This device has a series of integrated fluidic handling steps for the electrophoretic separation of dopamine from catechol at 2.2 and 4.6 mM, respectively. The technological advancement of this device consisted of miniaturized circuitry for both power and control of the ECD as mentioned previously [167]. As such, this device could be completely portable. Skelley et al. have developed a microfluidic device for the analysis of amino acid biomarkers for extraterrestrial analysis on Mars. This device integrated 34 valves and 8 pumps for the pneumatic control of fluids on-chip. The off-chip components include: pneumatic airlines, LIF detection equipment and the electronic/power control. All of the off-chip components were contained in a portable module which was self contained. This microfluidic device and portable module was capable of performing amino acid separations of seven amino acids in under 120 s with detection limits of 130 pM [159]. This work illustrates an exciting example of portable and remote analysis that does not rely on fully integrated systems. Culbertson et al. [168] have developed a portable microfluidic platform capable of separating a series of amino acids in a variety of reduced and hyper gravity environments. The microfluidic device integrates a series of sample handling techniques including electroosmotic pumps and non-mechanical gating followed by capillary electrophoresis. Portable off-chip components include high voltage power supplies, LIF detection equipment, accelerometer and electronics/power control and is illustrated in Fig. 9. This device was demonstrated to perform separations of a variety of amino acids at reduced gravity in under 12 s with <1% RSD. This device also has the footprint and payload limits that enable it to be transported by an unmanned rocket for extraterrestrial analysis of amino acids from a bioreactor. Cellar et al. [11,158] has demonstrated a highly integrated microfluidic device for the separation of neurotransmitters and modulators that were sampled from a push–pull probe. This device-integrated sampling from an external push–pull probe, mixing, derivatization, injection, and separation. The off-chip components of this device include: pneumatic actuation, LIF detection components and electronic/power control. This device separated seven neurotransmitters in under 20 s with detection limits of 60 nM and efficiencies of 100,000 theoretical plates.

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Fig. 9. Illustrates a microfluidic device coupled to a portable modular detection and operation apparatus. The modular components include: (1) high voltage power supply, (2) reed relays, (3) accelerometer, (4) LIF detection components, and (5) data acquisition. This device was capable of monitoring amino acids with high time resolution. [168] – Reproduced by permission from the American Chemical Society.

These devices have since been integrated with fiber optics and electrodes for both on-chip chemical detection and flow meters, respectively. Completely integrated and automated microfluidic platforms offer operational simplicity that enables these technologies to be transferred from an analytical chemistry laboratory to a neuroscience laboratory where limited understanding exists of analytical techniques, but where advanced neuroscience technology must be employed. Minas et al. used a CMOS process to integrate a spectrophotometer with a microfluidic manifold shown in Fig. 10 [88]. The microfluidic device had an integrated Fabry–Perot optical resonator and a detection and readout system both fabricated using CMOS processing. The quantitative analysis using UVvis absorption detection was employed to measure the uric acid and total protein content in urine samples. The fluid handling components of this device were performed off-chip. Choudhury et al. fabricated a microfluidic device containing an OLED and a luminescent sensing element for the detection of glucose [55]. The detection of glucose was performed in the range from 0.5 to 5 mg/mL. The miniaturization of such chemical processing can be envisioned to be useful in a number of point-of-care applications that aim at investigating metabolic diseases. 4.4. Applications Review of the assay systems described above suggest that applications for integrated microfluidic devices fall under two categories: (1) those that require portability and moderate footprints that can range from shoebox to suitcase sizes, and (2) those that require autonomous control and miniaturized footprints that are on the scale of a monolithic microfluidic device. We refer to each category as modular plug-in applications and autonomous miniature applications. The majority of the work in

Fig. 10. Integrated spectrophotometer on a microfluidic manifold. The spectrophotometer was fabricated using CMOS processing to integrate a light source, filter and photodiode. The spectrophotometer was bonded to a microfluidic manifold. The microfluidics served to mix analyte with colorimetric reagent that was consequently transported to the detection window for analysis. Uric acid and total protein was analyzed using this partially integrated device. Reproduced by permission of The Royal Society of Chemistry and the authors [88].

academia and industry has focused on developing microfluidic technologies for modular plug-in applications. Nevertheless, a trend toward devices that are autonomous monolithic systems is emerging and will enable more futuristic applications. Biochemical applications that are currently being pursued by modular plug-in microfluidic devices include: point-of-care diagnostics, space exploration, forensics, and counter terrorism. Point-of-care diagnostics can benefit from microfluidic technologies in that they provide cost-to-scale advantages in addition to improved throughput and efficiencies. This could provide methods to improve the inefficiencies and inherent costs that are often associated with healthcare testing in central laboratories. Furthermore, these devices can be envisioned to be fully portable and commercially available for home use which could provide the capability of daily health measurements. Such devices could also revolutionize detection in resource-limited environments such as Third World countries. Point-of-care devices can ideally be configured for a number of chemical tests including: blood chemistry, DNA sequencing, genotyping for forensic applications, biomarker analysis, breath analysis, and pathogen detection. In addition to cost-to-scale and efficiency advantages, microfluidic devices also offer the capability to automate parallel analysis. Their potential of disposability offers limited carryover for consecutive samples. Applications in space exploration include the development of microfluidic platforms for biosentinels or extraterrestrial biochemical analysis. Microfluidic devices are useful in space exploration due to their small lateral dimensions and hence reduced payload, reduced reagent and sample consumption which can typically be less than 1 mL per year, and reduced waste production. Overall, these devices can be used in environments where portable, rugged, and reli-

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able devices capable of sustained remote automated operation are required. Applications for autonomous microfluidic devices include a variety of health monitoring systems. Implantable microfluidic devices for biochemical analysis could provide real-time feedback on a variety of biological markers of clinical relevance in cancer, metabolic disease, microbial infection or internal organ failure. These devices could provide a method for early detection of these diseases or dysfunctions and hence provide better prognoses for treatment. These devices offer not only the capability of integration into a biological system where size requirements and autonomous control are critical, but could also be envisioned as being integrated into consumer products or defense systems. An example of integrated microfluidic devices for defense systems includes insect cyborgs which provide controlled transport for miniaturized analytical platforms. Overall, the miniaturization and consequently the size to scale advantages of such devices will further provide an amenable outlet to consumer, military or health care use. 5. Conclusions The tremendous research into separation systems that utilize fluidic manifolds mated with conventional power supplies, pumps, and detectors, over the past 14 years has generated a significant foundation of knowledge for microfluidic fabrication, chemical separation, and detection of biochemicals. From this foundation, a number of novel completely integrated microfluidic devices can begin to be realized and applied in a variety of biochemical applications where temporal resolution, spatial resolution, sample throughput, complex sample processing and/or small sample volume elements are required. Furthermore, integrated devices can literally go where non-integrated devices can not, whether they be implanted in the subcutaneous space of patients, embedded next to the microphone of a cellular phone for breath analysis, or inserted prior to surgery for the real-time analysis of cardiac markers. These devices truly have the capability to revolutionize the role of microfluidic devices and could also revolutionize the way we think about preventative and surgical medicine. For example, these devices could provide the technology to enable an individual to monitor their health in a routine way. There is a large gap between current instruments and these possibilities, but the first steps in this process have been taken with great success as documented by this review. The future in this field is likely going to be exciting as microfluidic technology continues to progress towards fully integrated devices for biochemical analysis. References [1] P.J.A. Kenis, R.F. Ismagilov, G.M. Whitesides, Science 285 (1999) 83. [2] A. Manz, N. Graber, H.M. Widmer, Sens. Actuators B 1 (1990) 244. [3] S.C. Jacobson, C.T. Culbertson, J.E. Daler, J.M. Ramsey, Anal. Chem. 70 (1998) 3476. [4] J.S. Jorgenson, K.D. Lukacs, Science 222 (1983) 266. [5] D.J. Harrison, K. Fluri, K. Seiler, Z. Fan, C.S. Effenhauser, A. Manz, Science 261 (1993) 895.

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