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Sensors and Actuators B: Chemical
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icrobead dispensing and ultralow DNA hybridization detection using liquidielectrophoresis
avi Prakash, Karan V.I.S. Kaler ∗
iosystems Research and Applications Group, Schulich School of Engineering, University of Calgary, 2500 University Drive NW, Calgary, Alberta T2N 1N4, Canada
r t i c l e i n f o
rticle history:eceived 31 January 2012eceived in revised form 26 April 2012ccepted 29 April 2012vailable online 4 May 2012
Chip based pathogen detection is one of the prime motivation for research in the field of microfluidics.Thus far, on-chip detection has primarily been demonstrated using closed-channel microfluidic devices,which are prone to certain limitations in terms of detection threshold and complexity of sample handling.Present day pathogen detection methods are often ‘bead based assays’ and their success is substantiallyinfluenced by the dispensing and manipulation capability of the utilized microfluidic technology. Thiswork utilizes liquid dielectrophoresis (LDEP) based surface microfluidics (SMF) devices to automaticallydispense ultra-low concentration of various microbead samples (different in size and functionalities)by judiciously adjusting conductivity and rheological properties of actuated carrier fluids. Control over
ead dispensingNA analysis
size and concentration of dispensed beads along with further manipulation and mixing for bead basedbio-assays has been successfully demonstrated. The application for the proposed rapid, ultra-low, bead-based DNA hybridization detection scheme, where 1–2 micro-beads can be manipulated in nanoliter sizedaliquots for detection in range of ∼100 attomoles, is envisioned for post amplification screening assays.In conclusion, we propose that LDEP based devices can be usefully applied and serve as alternative to theconventional closed channel device for ultra-low bead based detection and screening assays.
The domain of microfluidics has demonstrated considerablerogress towards the development of micro total analysis systemsMicroTAS) or, lab on a chip (LOC) devices [1], that integratesne or more of: sample handling, manipulation/mixing andetection/diagnostics on a miniaturized platform. These minia-urized devices require very small quantities of costly and oftenazardous bio-samples and with much shorter reaction timeso produce rapid screening and prompt diagnostics, renderinghem both portable and cost effective. Microfluidic devices havevolved from very simple flow-based schemes [2] to complexntegrated systems that have demonstrated applicability in severaliological procedures and assays including but not limited tolectrophoresis [3], polymerase chain reaction (PCR) [4], DNAnalysis [5,6], immunoassays [7], micro-organism detection [8]nd mimicking/analyzing cellular behavior [9]. Several researchroups have reported progress towards the development of smart,
n-chip bio-sensors, mostly using closed channel microfluidicevices [10–12]. In such devices, pathogenic organisms are oftenetected through their specific genomic RNA, based on the specific
DNA–RNA interactions [12]. Depending on the nomenclature ofthe pathogenic RNA, tailored oligonucleotide sequence can begenerated and used to detect the presence of specific pathogens ina bio-sample. The most commonly used detection scheme is a ‘beadbased’ assay, where the targeted DNA/RNA (from the pathogen)is covalently attached to a microbead through a linker moleculeand mixed with a complementary probe sequence (synthesizedbased on the knowledge about the target DNA/RNA). Suitablefluorophores/quenchers are incorporated during the hybridizationassay and the photo response following the hybridization can becaptured and quantified using Forster Resonance Energy Transfer(FRET) [13,14]. Bead based assays such as Luminex xMAP® arecommonly used in off-chip set-ups for post amplification screen-ing assays where the amplified target DNA is attached selectivelyto anchored markers and detected using flow cytometer [15].Alternatively, Surface Plasmon Resonance (SPR) and other opticalinterferometry techniques have been demonstrated for label-free detection of pathogens by utilizing evanescence waves atpatterned metal/dielectric interfaces [16]. Such schemes utilizenano-structured surfaces, functionalized with specific bio-markerswhich bind and detect the presence of specific pathogenic agent
strains by means of changes in the refractive index and associatedoptical and evanescent properties. Although such schemes havevalidated the concept of pathogen detection using a microfluidicdevice, they remain prone to drawbacks, such as extensive off-chip
ample preparation, requirement of multiple off-chip componentso generate uniform flow of anchored beads, and large amountf dead sample volumes. Due to such limitations, these schemesan only manipulate and detect a relatively high concentration ofargeted beads and higher concentrations of pathogen samples.
In this paper, we report on a liquid-dielectrophoresis (L-DEP)ased surface microfluidic (SMF) dispensing and manipulationevice that can rapidly dispense ultralow quantities of function-lized microbeads in multitude of low volume (nL to pL range)roplets and conduct bead based bio-assays. We have furthermoreemonstrated that such dispensing schemes can be tailored to dis-ense a wide range of microbeads in a very reliable fashion, thusaking them suitable and providing an alternative technology for
ead-based multiplexed high throughput screening and on-chipio-detection assays.
. Theory
The term dielectrophoresis (DEP) [17] is attributed to anylectro kinetic phenomenon which manifests itself on a dielec-ric particle (particle-DEP) [18,19] or in a dielectric liquid mediaLDEP) [20], when subjected to spatially non-uniform electric field.
hen applied to liquids, dielectrophoresis has been successfullyeveraged using surface microfluidic devices to dispense arrays ofltrafine, homogeneous [20,21] and/or, heterogeneous phase [22],niform or, variable volume droplets [23], in the volume rangef few nanoliter to picoliter. A number of studies have inves-igated the capabilities of LDEP actuation scheme in dispensingiosamples and macromolecules (such as lipids, quantum dots androteins) [24,25]. It has been shown that using dielectric mediauch as de-ionized (DI) water, which has fairly low conductivity�m ∼ 1 �S/cm), LDEP actuation scheme may be used for size basedeparation of micro-sized polystyrene beads. Jones et al. [19] haveeported that when polystyrene micro-beads (dia. ∼1 �m) are actu-ted within de-ionized water sample over a pair of uniform L-DEPlectrodes, their distribution is largely influenced by a positive DEPorce on the beads, subjected to regions of strong non-uniform elec-ric field, affecting their effective transport during the actuation.everal other DEP based electrode structures have also been uti-ized to fractionate microbeads and even living/dead cells usingither positive or negative DEP regimes [26]. The nature and effectsf the DEP force can be significantly altered by controlling the con-uctivity, dielectric properties of the fluidic medium, along with itsheological properties.
Now, in order to achieve ‘bead-based’ pathogen detection andcreening assays, LDEP devices should be capable of dispensingltralow quantities of biosamples and micro-beads; either uni-ormly or of controlled volumes, across the array of dispensedroplets. The experimental work reported in this paper wasonducted under specific fluidic properties to implement LDEPased dispensing as a tool for rapid screening of functionalizedicrobeads, in order to conduct specific DNA assay. The theo-
etical portion of the work focuses on optimization of the fluidicedium properties in order to ensure that DEP forces exerted on
he microbeads do not adversely impact their reliable dispensing,s required by targeted bio-applications.
During LDEP actuation of the liquid jet, the suspended particles,re swept along the jet flow direction are subjected to transverseorthogonal to the flow) DEP force, as a result of the non-uniformlectric field emanating from the co-planar parallel electrodeshen biased by A.C. voltage. The DEP force acting on a spherical
haped particle can be expressed as [19]
�DEP = 2�εmR3 Re[K] �∇
(∣∣Erms∣∣2
)(1)
ctuators B 169 (2012) 274– 283 275
where R is the particle radius, εm is the medium permittivity, E isthe non-uniform electric field and K is the frequency and mate-rial/media property dependent Clausius–Mossotti factor
K(ω) = εp − εm
εp + 2εm, ε = ε − j
�
ω(2)
The quantity ε represents the complex and frequency depen-dent permittivity of the particle, and that of the liquid media, ωis the frequency of the applied electric field and � is the electricalconductivity. The sign of Re(K) determines the direction of the DEPforce and its value lies in the range [−0.5 to 1.0], in different regionsof the frequency spectrum. The four parameters: εp, �p, εm, �m, col-lectively modulate the DEP force, resulting in positive or negativeDEP of particles in different regions of the frequency spectrum.
The cross-over frequency(ωc) is a threshold frequency at whichthe DEP force switches sign, either from positive to negative orvice-a-versa, given as [27]
ωc =√∣∣(�m − �p)
∣∣ (�p + 2�m)(∣∣εp − εm∣∣) (εp + 2εm)
(3)
Thus based on this simplified particle model, there are tworegimes of DEP forces separated by the cross-over frequency (ωc).The Clausius–Mossotti factor, for frequencies very low or very highas compared to the corresponding ωc, is dependent on the con-ductivity and permittivity differences between the particle andmedium, respectively:
At low frequencies (ω << ωc) : K0 = �p − �m
�p + 2�m(4)
At high frequencies (ω >> ωc) : K∞ = εp − εm
εp + 2εm(5)
The electrical conductivity of the micro-particle is the mostcumbersome to evaluate, as it depends on both the bulk con-ductivity (�b) and the surface conductance of the particle whoserespective contributions are as follows:
�p = �b + 2(Ks + Ki)R
(6)
where Ks is the surface conductance of the microbead and Ki is theconductance due to the ionic, diffused double layer. In case of latexbeads with a polystyrene core and shell, the bulk conductivity isnegligibly small (�b ∼ 0) whereas, Ks ∼ 1.2 nS (referred from [28])and in very low ionic media (such as glycerol, very low conc. saltsoln.), term Ki can be neglected when modeling the particle DEPresponse.
The DEP force under the assumption that non-uniform electricfield inside the aqueous jet is nearly azimuthal in form
E�(r) ≈ Va
�r(�) (7)
Since the electric field is capacitively coupled, the voltage acrossthe liquid finger is represented by
Vf ≈ Cd
2Cm + CdVa (8)
where Va is the applied ac voltage. The expressions for the capaci-tance terms, for both uniform and tapered liquid jet actuation, havepreviously been reported in [20,22] and are related to the geomet-rical and material properties of liquid actuation electrode structure.Using Eqs. (1) and (7), the DEP force can be determined [19]
FDEP,r = −4εmR3 Re [K]V2
(r)
(9)
�r3 f
To achieve a uniform microbead dispensing across the dropletarray we need to minimize the impact of the sedimentation andjudiciously tailor the DEP force on the microbeads during L-DEP
276 R. Prakash, K.V.I.S. Kaler / Sensors and Actuators B 169 (2012) 274– 283
Table 1Rheological and dielectric properties of DI water–glycerol solution.
Table 2Properties of various microbeads used in experiments.
Bead type Notation Diameter (�m) �Ex./Em. (nm)
Green 1 G1 1.2 450/540Dynabead® (functionalized) DB 2.8 450/540Red 1 R1 5.7 540/680
(DNAconc.)before
Sample 6 50 1250 63 16.0
ctuation. To achieve this, we selected fluidic media whose density,iscosity, conductivity and dielectric constant maybe be suitablyltered and tailored to achieve the uniform dispensing require-ent.In order to avoid the induction of positive DEP forces, at 100 kHz
LDEP actuation freq.), it is required that the media conductivitye suitably adjusted such that �m > �p; and the DEP force is neg-tive thus pushing the microbeads upwards and away from theigh field region located at the electrode surfaces. We have usedomogeneous DI–glycerol mixtures to alter the rheological andlectrical properties of the fluid media (Table 1). Increasing thelycerol concentration in the aqueous media results in an increasen density and viscosity while the dielectric constant of the mixtureecreases with a controlled increase in the medium conductivityhich stays in the range of (2.5–15 �S/cm). Also, glycerol is a suit-
ble media to handle bio-samples such as DNA/RNA as the osmoticressure of such mixtures are closely comparable to that of cellularuids and thus biocompatible. Low concentration ionic solutionsnd non-ionic buffers (TRIS(tris(hydroxymethyl)aminomethane)-ES(2-(N-morpholino)ethanesulfonic acid)) were used to prepare
nd actuate the functionalized microbeads in an optimized mediumonductivity, ranging from 10 to 15 �S/cm.
Numerical analysis was conducted to determine the range ofariance of FDEP and cross-over frequency for the different flu-dic media (Table 1) and bead radii, under the conditions of LDEPctuation (AC frequency = 100 kHz and actuation voltage = 450 V).ig. 1 indicates the nature of DEP force experienced by the beadsf diameter range 2–15 �m, suspended in different conductivityedia, suggesting that medium conductivity of nearly 8–12 �S/cm
s required in order to avoid the regime of positive DEP. The vari-tion of cross-over frequency was also analyzed in order to ensurehat the LDEP actuation frequency (100 kHz) is lower than the cross-ver frequency for the entire experimental regime (data points inig. 2), hence justifying the use of Eq. (4) in the numerical model.igs. 1 and 2 confirm that the effective force acting on the particlesuring such LDEP actuations will result in an upward trajectory forhe microbeads within the transient flow. It does however appearhat particles with diameters less than 1 �m will still be subjectedo a significant positive DEP force. Although for particles of thisize, the positive DEP force (∝R3) based attraction of beads towardshe chip surface is largely restricted by the forces attributed dueo Brownian motion (∝R−1), which is found to overpower theositive DEP force and maintain a uniform dispersion of theseicrobeads [26]. The numerical analysis was further validated dur-
ng the experimentation where the fluidic media properties wereailored to achieve the uniform flow and dispensing of the different
icrobead samples during both uniform and variable volume LDEPispensing. Results and further validation of the numerical analysisre presented in Section 4.
The theoretical analysis conducted in this work confirms
hat under specific fluidic conditions, LDEP based actuationan be utilized for controlled dispensing of microbeads, withpplications intended for chip based bio-assays. The analysis how-ver is restricted to simple microbead models and will require
improvement to analyze bead surfaces that are functionalizedwith DNA/RNA and other linker molecules. Our experiments haveso far suggested that even functionalized microbeads, with bio-molecules bound to bead surface, can be very uniformly distributedduring the LDEP based dispensing scheme as demonstrated in thebead based DNA hybridization assay, in Section 4.
3. Materials and methods
3.1. Device fabrication
The micro-fabricated SMF device, used for experimentation,consists of a pair of patterned co-planar metal electrodes coatedwith a thin insulating dielectric layers, deposited over a passivatedsilicon wafer (SiO2 on Si). In this investigation, we implement SMFchips that consist of several uniform and continuous tapered L-DEPelectrode schemes and an integrated scheme of two LDEP and DDEPelectrode structures [21], all shown in Fig. 3. To fabricate these SMFdevices, initially a 150–200 nm thick metal (Al or ITO) layer wasdeposited over a passivated silicon substrate, using a magnetronsputtering process and then patterned using standard lithogra-phy and wet etching procedures. The silicon wafers, housing thepatterned electrodes, were then passivated with a thin insulatingdielectric (450 nm of Si3N4) using plasma enhanced chemical vapordeposition (PECVD) method and subsequently patterned using areactive ion etching (RIE) process. The top surface of the SMF devicewas rendered hydrophobic by spin coating a thin layer (∼100 nm)of Teflon® AF 2400 (Grade: 400S1-100-1, DuPont Inc., USA) to facil-itate droplet dispensing and further actuation.
3.2. Sample preparation
Five different bead samples of various sizes and surface func-tionalization tags (see Table 2 and Fig. 4) were used to demonstratethe handling and dispensing capabilities of LDEP actuation underthe controlled fluidic conditions. Three different ssDNA oligonu-cleotides (Integrated DNA Technologies (IDT), USA) were preparedin 10–15 �S/cm TRIS–MES buffer (pH = 7.8) and utilized in the on-chip, ‘bead-based’ hybridization assay.
The structural and other attributes of these four ssDNA samplesare shown in Table 3 and Fig. 4. In order to perform quantitativeassays, the concentration of all bio-samples was measured dur-ing sample preparation using Nanodrop 1000 spectrophotometer(Thermo Scientific, USA) and further during bio-detection assay,dispensed and mixed bio-samples were investigated using a PhotoMultiplier Tube (PMT).
The binding efficiency (BE) of dynabead® (DB) for the 20 bpoligonucleotides was obtained during the coupling of ssDNA (S1) toDB, by comparing molecular concentration before and after off-chipstreptavidin–biotin binding, using the expression:
BE = (DNAconc.)before − (DNAconc.)after × 100 (10)
where (DNAconc.)before is conc. of oligonucleotide in the samplebefore binding onto the DB and (DNAconc.)after is final conc. ofssDNA molecules remained in the bio-sample after binding onto
R. Prakash, K.V.I.S. Kaler / Sensors and Actuators B 169 (2012) 274– 283 277
+ve D
EP
re
gim
e-v
e D
EP
re
gim
e
Bead Radius (μm)
1 2 3 4 5 6 7 8
FD
EP (
N)
-80 x10-9
-60 x10-9
-40 x10-9
-20 x10-9
0
20x10-9
40x10-9
60x10-9
80x10-9
σm = 1 μS/cm
σm = 15 μS/cm
Fig. 1. Plots showing variation of DEP force for different sized microbeads actuated in varying conductivity fluidic media at 100 kHz.
Bead Radius (μm)
1 2 3 4 5 6 7
Cro
ss-o
ver
frequ
ency (
MH
z)
0.0
0.5
1.0
1.5
2.0
2.5
3.0
σm = 15μS/ cm
σm = 1 μS/ cm
fact = 100 kHz
or diff
telosbcts
TA
Fig. 2. Simulation result shows variation of cross-over frequency f
he DB, during the off-chip functionalization of dynabeads. Bindingfficiency data calculated using this empirical method was uti-ized to evaluate the quantity of oligonucleotide markers anchoredn the microbeads which allowed us to calculate the amount ofsDNA sample quenched and successfully detected during the bead
ased hybridization assay, as reported in Section 4. For the oligonu-leotide samples used in this work, BE was found to range from 75%o 80% during multiple measurements and the amount of boundsDNA per bead was found to be ∼60 attomoles.
able 3ttributes of various bio-probe molecules used in experiments.
erent microbeads, actuated in different conductivity fluidic media.
Quenching efficiency (QE) for the FRET/Quenching, during thebio-assay was evaluated using the PMT Photocurrent (Ip) and anempirical relation given by Eq. (11):(
(Ip)unmixed − (Ip)mixed
)
QE =
(Ip)unmixed − (Ip)bck× 100 (11)
where (Ip)unmixed is the PMT photocurrent (Ip) measured beforemixing, (Ip)mixed is measured after binary mixing is performed
278 R. Prakash, K.V.I.S. Kaler / Sensors and Actuators B 169 (2012) 274– 283
Fig. 3. (a) Uniform LDEP electrode structure and its attributes; (b) continuous tapered, pinched LDEP electrode structure (both with and without bumps) and (c) a simpleintegrated, 1 × 1, LDEP–DDEP electrode scheme for assays.
Fig. 4. Stoichiometric details of the oligonucleotide samples and schematic of thebead based FRET assay.
during the assay and (Ip)bck is the background signal. The quenchingefficiency (QE) data along with the calculated fluorophore concen-tration and the known binding mechanism of streptavidin–biotin(binding ratio for used ssDNA molecules: 1:2.5), allows quantifi-cation of hybridized oligonucleotide probes detected during theDNA hybridization assay and hence determination of the limit ofdetection (LOD) for the assay. The merit of this bead-based DNAhybridization is the FRET based quenching, illustrated earlier inFig. 4 and the LOD reported in this work refers to the thresholdoligonucleotide concentration that was successfully detected usingthe PMT. The detection threshold is based on the PMT photocurrentand the QE reading for the lowest ssDNA concentration which canbe resolved above the background and the non-specific nucleotidequenching due to adjacent nucleotide (∼10–20%), which is report-edly very low as compared to FRET based/static quenching [29].
3.3. Experimental procedures
A schematic diagram of the optoelectronic setup used for exper-imentation has been reported in [21]. The SMF chip was securedusing spring loaded pogo pins onto a PCB to provide electricalconnections to the bonding pads. The arrangement was placed on
a fluorescent microscope platform (BX51, Olympus, Japan) cou-pled to a photomultiplier tube (PMT) (H-7468-10, Hamamatsu,Japan, operated at PMT gain of 106) for measuring the fluorescenceintensity and a high speed CMOS imager (Mega speed) or, a CCD
s and A
iatarTor2diwa
ccmas
4
4
bdmhmaaeTntLt1
Fmi(i
R. Prakash, K.V.I.S. Kaler / Sensor
mager (QImaging, Canada) to observe the dispensed biosamplesnd capture images of the conducted assays. A signal genera-or (TGA1244, TTi, UK) and a high-voltage, high-frequency powermplifier (Precision Power Amplifier 5205A, Fluke) provided theequired AC voltage to drive the coplaner electrode arrangement.he actuation process was controlled using a software driver devel-ped using labview (NI Labview, USA) and the output data wasecorded either in form of high speed videos (original frame rates:000–2500 fps) using the high speed camera or, images and PMTatasets for quantitative assay. The high speed videos were dig-
tized and an image probing program (provided by Mega speed)as used to capture the dynamics of microbeads during LDEP
ctuation.The LDEP actuation on electrode schemes shown in Fig. 3 was
onducted by energizing the electrode pair using AC voltage, typi-ally, 400–500 Vpp at a frequency of 100 kHz. Droplet transport andixing was achieved by applying a much lower voltage and low
ctuation frequency (typically 100 Vpp @ 30–50 Hz) to the electro-tatic droplet transport scheme (see Fig. 3c).
. Results and discussion
.1. Uniform dispensing of low concentration latex microbeads
In Section 2, numerical analysis was conducted to investigate theehavior of the five different microbead samples (Table 2), actuateduring this work. It was established that by judiciously choosingedium conductivity (in range of 8–15 �S/cm) and sufficiently
igh medium density and viscosity, controlled dispensing of theseicrobead was plausible. Under these conditions, microbeads stay
float, close to the aqueous-oil interface during LDEP actuationnd are carried away by the actuated liquid jet. The fluidic prop-rties of various media used during experiments are reported inable 1. Based on the experimental observation, along with theumerical model, it was established that Glycerol–DI water mix-
ure with Glycerol concentration of 30–40% by vol. was suitable forDEP actuation of microbeads using both uniform and continuousapered electrode schemes. Each microbead sample was actuated0–15 times using uniform LDEP electrodes (w = g = 20 �m, 30 �m,
ig. 5. Micrograph collage showing different concentrations and sizes of microbeads acicrobead (NF2: dia. – 15 �m) during LDEP actuation on w = g = 20 �m LDEP electrodes (s
mage of a dispensed droplet (∼1 nL) with large amount of (100 beads/nL) G1 beads; (h, iparent conc.: 4–5 × 106 beads/mL); (j) a single NF1 bead dispensed in ∼300 pL droplet durmages of three adjacent droplets (∼1 nL) containing pair of R2 beads during LDEP actuat
ctuators B 169 (2012) 274– 283 279
shown in Fig. 3a) and the bead concentration was measured inthe array of dispensed uniform volume daughter droplets. Further-more, the bead concentration in daughter droplets was normalizedand compared with its concentration in each parent droplet in orderto overcome effects of variations due to off-chip sample preparationsteps. The resulting distribution of dispensed bead concentrationsamong the daughter droplets, for different bead sizes and parentconc., is shown by a collage of micrographs in Fig. 5. The variationin normalized bead concentration during these LDEP actuations isdepicted in Fig. 6.
Results reported in Figs. 5 and 6 suggest that the LDEP actuationsin the tailored Glycerol–DI water medium resulted in a highly uni-form dispensing of the various bead samples, ranging in diameterfrom 1 �m to 15 �m, in the midsection of the electrode structure,corresponding to droplet 2–6 (see plot in Fig. 6). In order to avoidcoagulation and steric-hindrance for the larger size beads (dia.10–15 �m), where bead diameter became comparable to the liquidjet radius, very low concentrations were used in the experiments(Fig. 5h–l).
The results shown in Fig. 6 also reported an excess dispensingor, accumulation in the end droplets for all microbead actuationswhich can be explained using the underlying fluid dynamics. Dur-ing an LDEP actuation, the liquid jet is propelled very rapidly(actuation speed ∼20 cm/s) and at such high speeds, the viscousdrag forces acting on the particle oppose the forward propagationof the microbeads with the moving jet. This intermittent viscousdrag force results in comparably slower moving microbeads in thepropagating jet. The inertia of these microbeads carries them for awhile through the jet, even after the jet propagation has stopped.So, in order to allow uniform dispensing throughout the dropletarray, the liquid jet was allowed to stay formed longer than usualto ensure that beads have saturated in the flow. As a result of theseflow irregularities, the end droplets accumulate a relatively higherconcentration of microbead samples. This phenomenon does notaffect the reliability of the dispensing scheme as the end droplets
can be considered and opted out as the dead volume during therapid dispensing mechanism and furthermore they establish thefact that positive DEP based pinning does not take place at all inthe specific fluidic samples chosen during the LDEP actuations. The
tuated and dispensed using LDEP. (a–e) Transport and dispensing of polystyreneee Video 1 – Uniform LDEP bead dispensing NF1); (f, g) bright field and fluorescent) two adjacent ∼0.8 nL daughter droplets with 3 NF2 beads during LDEP actuationing actuation on a w = g = 10 �m electrode scheme; (k, l) bright field and fluorescention process.
280 R. Prakash, K.V.I.S. Kaler / Sensors and A
FLb
aigffiajotsd
4v
paacpuadgiLicswaFadcdF1e
s
ig. 6. Plots show distribution of various beads during LDEP actuation on a uniformDEP electrode scheme (w = g = 20 �m), over 10–15 repeated actuations for eachead sample.
ctuation time shots shown in Fig. 5(a)–(e) also show that dur-ng the jet actuation, the beads appear to move precisely along theap between the LDEP electrodes and near the aqueous–oil inter-ace of the hemi-cylindrical jet profile, while being farthest awayrom the high field regions of the two electrodes. The very highnterfacial energy required to push such large beads through thequeous–oil interface ensures that the beads remain in the aqueouset during dispensing. This feature of LDEP actuation has facilitatedther potential particle dispensing applications, such as deposi-ion of lipid membrane on the actuated microbead/nanoparticleurface during an emulsion jet actuation with lipid-in-mineral oilispersion [24,25].
.2. Controlled dispensing of various sized latex microbeads usingariable volume dispensing scheme
Ideally, an on-chip dispensing scheme should be able to dis-ense controlled volume and amounts of different bio-samplesnd macromolecules. In previous work, we have reported on suchn LDEP electrode structure, with a continuous tapered design,ontaining specifically positioned pinches that allow controlled dis-ensing and positioning of variable volume sub-nanoliter dropletssing LDEP [23]. In order to establish the SMF technology is a suit-ble mean for bead based bio-assay applications it is important toemonstrate that the control over the dispensed mass/sample conc.radient can be reliably achieved for the range of micro-beads usedn on-chip pathogen detection applications. The variable volumeDEP dispensing scheme utilized for these experiments is reportedn Fig. 3(b). Microbead samples of different bead sizes and parentoncentration were actuated on the continuous tapered electrodecheme. The actuation resulted in dispensing of daughter dropletsith volume ranging from 1.9 nL to 300 pL. The microbeads were
gain actuated in Glycerol–DI water media (30–40% by vol., seeig. 7a–f) and the dispensed bead concentrations were measurednd averaged over at least 10 LDEP actuations per bead sample. Theistribution of dispensed microbeads was found to be very reliablyontrolled by the electrode structure, resulting in a bead count gra-ient, in the dispensed droplet arrays as shown in Fig. 7(g)–(i) andig. 8. Here the number of dispensed beads is reduced from 4 to
with controlled reduction in droplet volume over the taperedlectrode scheme.
The plots in Fig. 8 also show that although the taperedcheme allows user to facilitate dispensing of varying number of
ctuators B 169 (2012) 274– 283
microbeads in the dispensed droplets, it results in a very uniformbead concentration distribution, comparable to the parent dropletconcentration. Such schemes are potentially useful as they can trimthe number of functionalized beads to very low quantities (evenachieving 1 bead per droplet) because of the order of magnitudevariation in the dispensed droplet volumes. Variable volume beaddispensing schemes can also be leveraged in integrated bio-assaysto observe the detection thresholds and limiting marker concen-tration during a chip based bioassay.
The experimental results reported in these two sections suggestthat LDEP is a reliable scheme for the rapid and controlled dis-pensing of ultrafine micro-bead and bio-sample quantities whichcan then be further utilized downstream to conduct on chip bio-assays with ultralow samples. The next section implements theLDEP based uniform bead dispensing scheme to demonstrate afundamental quantitative bead based DNA hybridization assay.
4.3. Bead based assay utilizing selective oligonucleotidehybridization and quenching
The dynabead® sample (DB) used for this experiment is reportedin Table 2. These microbeads contain an outer shell of streptavidinlinkers which readily bind to biotin functional group. A suitableoligonucleotide sample, S1 with a biotinylated 5′-end and a flu-orophore at the 3′-end was bound to the beads to prepare thebio-probe (see Fig. 4). Detailed attributes of the binding process arereported in Section 3, leading to preparation of the dynabead® sam-ple in Glycerol–DI water mixture (30% by vol. Glycerol; 0.005% byvol. Tween-20). A complementary oligo sample P1 with a quenchermolecule attached to its 5′ end (see Section 3) was also preparedin identical aqueous media (concentration: 50 �g/mL). A simple1 × 1 electrode scheme, shown in Fig. 3(c), was used in the bio-assay. Firstly, 1 �L parent sample droplets of the functionalizedmicrobead sample (bead-S1) and the complementary oligo sam-ple (P1) were mounted to the two parent sites of the electrodescheme. Upon L-DEP actuation, two arrays of equivolume daughterdroplets were dispensed on the two LDEP electrodes. Subsequentlythe DDEP electrodes were energized and pair of droplets from thetwo arrays were mixed to facilitate interaction between the beadbased bio-probe and its marker sequence (Fig. 9a–d). The mixedand unmixed droplets were analyzed using the opto-electronic set-up in order to quantify and detect the selective hybridization anddetection of the specific DNA sequence (Fig. 9). PMT was used tomeasure the fluorescence from the DBs in the droplets dispensedover the LDEP electrodes (Ipunmixed − Ipbck) and in the mixed droplet(Ipmixed − Ipbck). The measured photocurrent data is shown in thecaption of Fig. 9. This PMT data was used to calculate the associatedquenching efficiency (QE) which quantifies the DNA hybridization.
Upon successful hybridization, the fluorophore attached to themicrobead is dominantly quenched which results in a reduced flu-orescence emission (QE ∼ 90%) from the microbeads (Fig. 9h andl). The difference in the fluorescence intensity (Fig. 9f, h and j, l)is quantified (Ipunmixed − Ipmixed) in order to measure the presenceof the complementary DNA molecules. The microbead concentra-tion was lowered (Fig. 9e–h) during these experiments in orderto measure the detection threshold which is limited by the signalloss due to the off-chip detection set-up and the signal-to-noise(Ipbck) ratio. Since the biosamples used here were fairly simple;assay and detection times were less than a minute. The limit ofdetection (LOD) for our bead based DNA hybridization assay hasbeen found to be close to ∼100–200 attomoles for the ultra-lowbead assay (Fig. 9e–h) where a minimum of 2–3 functionalized
beads were investigated with the PMT before and after mixing.The DNA hybridization was also quantified in case of high beadconcentration and the amount of hybridized DNA probe was foundto be ∼5–10 femtomoles. The LOD, which was controlled by the
R. Prakash, K.V.I.S. Kaler / Sensors and Actuators B 169 (2012) 274– 283 281
Fig. 7. Images of variable volume and controlled bead dispensing using tapered LDEP scheme. (a–f) Transport and dispensing of polystyrene microbeads (dia. = 15 �m) duringLDEP actuation on continuous tapered (from w = g = 40 �m to w = g = 10 �m) LDEP electrodes (see Video 2 – Continuous tapered LDEP variable bead dispensin NF1); (g–i)bright field image showing controlled reduction in number of beads (R2 beads) with tapered LDEP based variable volume droplet dispensing.
F ts, fols
oDdc[
TQ
ig. 8. Plots showing distribution of different beads in dispensed daughter droplecheme (Fig. 3b).
ff-chip streptavidin–biotin binding, the ultra-low functionalized
B dispensing in nL aliquots, mixing and the high gain PMT basedetection, is an improvement when compared to the conventionallosed channel DNA hybridization assays (LOD ∼ few picomoles)30,31] but is higher than the LOD achieved through RT-PCR based
able 4uantitative analysis of probe (P1) DNA hybridization with ultra-low bead aliquots.
P1 conc. (�g/mL) Ipunmixed (�A) Ipmixed (�A)
0.5 7.55 3.97
1 7.50 3.29
10 7.60 2.04
25 7.52 1.8
50 7.55 1.7
lowing more than 10 L-DEP actuations per bead sample on the tapered electrode
amplification assays in microfluidic DNA/RNA detection schemes
where numerous PCR cycles (>10 cycles) are often required toachieve ultra-low LOD [30,32]. In contrast, the reported value ofLOD was obtained within a minute of sample dispensing and calcu-lated using Eqs. (10) and (11) and the measured PMT and Nanodrop
282 R. Prakash, K.V.I.S. Kaler / Sensors and Actuators B 169 (2012) 274– 283
Fig. 9. Snapshots showing experimental results from on-chip bead based bio-assays. (a, b) Mixed and unmixed sample/reagent droplets; (c, d) unmixed daughter dropletswith functionalized dynabeads® and corresponding markers prior to and after binary mixing; (e–h) micrographs showing bead based assay achieved at ultra-low beadconc. (2–3 beads per droplet); (e, f) bright and fluorescent image of unmixed and unquenched functionalized microbeads (Ipunmixed − Ipbck = 8.6 × 10−6 A); (g, h) quenchedmicrobeads in mixed droplet (QE ∼ 91%) as a result of DNA hybridization (Ipmixed − Ipbck = 0.9 × 10−6 A); (i–l) similar bead based assay results for relatively larger bead conc.(∼10 beads/nL) with (Ipunmixed − Ipbck = 2.2 × 10−5 A), (Ipmixed − Ipbck = 2.6 × 10−6 A) and QE ∼ 88%; (m–p) results of a control experiment showing mixing but no quenching fornon-specific bio-marker at low bead concentration and 50× optical magnification (Ipmixe − Ip = 6.8 × 10−6 A) and QE ∼ 20–25% (m, n) and high bead concentration (o, p)( r illust
difabppPseDuFTrc
flotpalbtbwb
Ipmixed − Ipbck = 7.0 × 10−6 A) and QE ∼ 20–25%. The binary mixing process is furthe
ata. However, one restriction for this ultra-low bead based assays that it still requires a relatively high DNA conc. (femtomolar)or the parent droplet and the low detection limit is achieved as
direct result of conducting the DNA hybridization on very smallead count (2–3 beads) in the dispensed daughter droplets. Theroposed bead based DNA hybridization assay is most suited forost amplification screening assays where off-chip or on-chip RT-CR amplicons can be screened with labeled bead based targetequences, while using very small amounts (few microliters) ofxpensive PCR products and marker sequences. Different probeNA concentrations were investigated using the bead based assaysing the ultra-low bead concentration (2–3 beads/nL), shown inig. 9(e–j), and the quantitative results are reported in Table 4.he reduced QE corresponding to lower probe concentration is aesult of poor FRET based quenching at relatively low quencheroncentrations (Bead-bound S1 concentration ∼0.5 �g/mL).
A control assay was also conducted to validate that the loss ofuorescence intensity and the resulting QE were indeed as a resultf detection of hybridization of complementary DNA sequences. Inhe control experiment, a non-complementary marker oligo sam-le S2 (see Fig. 4) was used to functionalize the DB microbeads. Thessay was repeated and the mixed and unmixed droplets were ana-yzed using the PMT. The resultant mixed droplets for this assay, foroth low and high bead conc. (shown in Fig. 9m–p) suggest very lit-
le change in fluorescent intensity (QE < 20–25%) of the bead basedio-probes which confirms that the particular sequence of markeras not present in the sample oligonucleotide. Also, since in bead
ased assay, fluorophore labeled sample DNA is stably anchored
d bck
rated in the video file: Video 3 – 1 × 1 binary mixing assay.
to the bead surface, static or random quenching which is alwayspresent and inversely affects the LOD for hybridization assays indroplets or micro-channels, is minimized. This results in superiorcontrol over the experiments and further improvement in confi-dence of detection and LOD for the bead based assay.
Such a bead based nucleic acid assay is also considered supe-rior in terms of the assay cost to the conventional bead assaybecause the amount of functionalized microbeads can be sub-stantially reduced (up to 2–3 beads per droplets and up to 1000beads per �L as shown in Fig. 9e–h). Furthermore these ultrafinemicrobead samples can be better maneuvered and detected withease using simple on-chip or, off-chip optical set-ups. The beadbased assay is also superior to the simple droplet based detectionschemes, as previously reported in [21], based on the fact that theamount of microbeads can be more conveniently tailored to lowernumbers. Since the amount of bound DNA sample is controlled bythe number of binding sites and the binding efficiency (see Eq. (10))for the particular biotinylated ssDNA molecule, quantification ofDNA hybridization in the lower concentration range is far morereliable using a bead based assay. The LOD achieved in this beadbased assay is found to be fairly comparable to other on-chip beadbased hybridization detection assays reported in recent literature[33].
5. Conclusions
The reported work demonstrates the utility of LDEP electrodeschemes for precision dispensing of ultra-low amounts of both
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degree at the University of Wales in 1981 and joined the University of Calgary in
R. Prakash, K.V.I.S. Kaler / Sensor
unctionalized/non-functionalized beads in the diameter range of–15 �m. The rapid, ultralow bead dispensing scheme is then
mplemented in a bead based DNA hybridization detection assay,here a targeted ssDNA, functionalized on bead surface, was used
o detect and quantify complementary Probe sequence at very lowead concentration (2–3 beads/nL) using FRET based quenching.he performance of the assay scheme was measured and reporteduantitatively using the off-chip PMT based detection and LOD wasound to be ∼100 attomoles for different probe DNA concentrations.his off-chip detection scheme somewhat restricts the detectionhreshold which can be further lowered by using fiber-optics andhin film waveguide integration to illuminate and collect fluores-ent signal from the SMF chip. Multiplexing of the assay unit to
8 × 8 matrix, where the full potential of the rapid dispensingcheme can be utilized to conduct post amplification screeningssays, such as the Luminex® bead assay [15], is a prime futurearget towards further improving the applicability of the proposedMF devices.
cknowledgments
The authors would like to acknowledge the financial supportrovided by NSERC Canada for the research presented in this papernd CMC Microsystems Canada for their support in the fabricationf the surface microfluidic devices.
ppendix A. Supplementary data
Supplementary data associated with this article can be found, inhe online version, at http://dx.doi.org/10.1016/j.snb.2012.04.081.
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Biographies
Ravi Prakash was born in India in 1985. He earned his Bachelor’s degree in Mechan-ical Engineering, in 2008, from Indian Institute of Technology Madras, India andjoined University of Calgary in 2008. He completed his Master’s degree from the Uni-versity of Calgary in 2010 and is now a PhD candidate working with Dr. Karan Kaler.In the past four years, Ravi has conducted research in the field of surface microflu-idics and its application towards bio-detection. His present research focus is ondeveloping low cost, miniaturized devices for bio-diagnostics and high throughputscreening assays.
Karan V.I.S. Kaler, born in India in 1951, is a Professor of Electrical and ComputerEngineering. He is registered professional Engineer and an elected fellow of theAmerican Institute of Medical and Biological Engineering. He earned his doctoral
1982. Dr. Kaler’s research laboratory, over the last 30 years, has engaged in the devel-opment and application of biological dielectrophoresis. His present research focuseson the development of integrated chip based microfluidic platform technology, tar-geted to the detection and analysis of biological cells and biomolecules.
