Optical sensing for on-chip digital microfluidics Jacqueline Nichols, Emily L. Landry, Brandon Born, Michael Wiltshire, Christopher M. Collier, and

Jonathan F. Holzman

University of British Columbia, 3333 University Way, Kelowna, BC, Canada V1V1V7

ABSTRACT

A digital microfluidic architecture is introduced for micron-scale localized fluid actuation and in in-situ optical sensing. Contemporary device integration challenges related to localization and device scalability are overcome through the introduction of a bi-layered digital microfluidic multiplexer. Trinary inputs are applied through differential combinations of voltage signals between upper (column) electrodes and lower (row) electrodes. The ultimate layout provides increased scalability for massively parallel microfluidic actuation applications with a minimal number of inputs. The on-chip sensing technique employed here incorporates a microlens in a folded-cavity arrangement (fabricated by a new voltage-tuned polymer electro-dispensing technique). Such a geometry heightens the sensitivity between the optical probe and fluid refractive properties and allows the device to probe the refractive index of the internal fluid. This optical refractometry sensing technique is merged with the actuation capabilities of the digital microfluidic multiplexer on a single lab-on-a-chip device. Keywords: Digital microfluidics, lab-on-a-chip, electrowetting, refractometry, optical sensing

1. INTRODUCTION High-sensitivity microfluidic reactors have significant potential in biotechnology, environmental, and physiochemical applications.1,2 Integrated microsystems have decreased reagent volumes and analysis times due to improvements in throughput and sensitivity.3 Such integrated devices are often implemented as continuous flow microfluidic systems with micromachined channels, pumps, and valves for controlled fluid flow in application-specific lab-on-a-chip devices.4,5

Real-time adaptive processing is desirable for many contemporary analytical technologies, though such reconfigurability is not compatible with permanently-etched/machined features in continuous flow structures. Digital microfluidics, where discrete fluid is actuated on a programmable and reconfigurable substrate, has emerged as a response to this adaptability need.6,7,8 A two-dimensional (2-D) electrode plane can be used with independent and programmable voltage control initiating localized fluid actuation. The electrohydrodynamic (voltage) actuation process9 is inherently reversible and programmable for complete system reconfigurability. The 2-D layout is programmed for analytical tasks defined by the user and can be programmed for real-time fault detection, process scheduling, and path planning.10

Actuation scalability is a primary challenge in digital microfluidics. The 2-D structures demand fine spatial resolutions with numerous inputs for complete 2-D fluid actuation and control. The standard actuation technique utilizes M rows and N columns of electrodes in a 2-D square-electrode grid. As the size of the grid is increased to sizes greater than 2 × 4,11 the number of electrical address lines required for the input of the applied voltages becomes infeasible without overlapping, shorting, or crossing address lines. Multi-level topologies that incorporate electrical via holes12 can be used to alleviate the difficulties with scalability, although they are typically avoided due to photolithographic complexity and fabrication costs. To facilitate actuation, one must ease the constraints associated with electrical addressability and reduce system inputs for digital microfluidic architectures.

Real-time sensing is a secondary challenge in digital microfluidics. It is necessary to probe in real-time fluid state information with sufficient sensitivity. Signal levels of contemporary sensing techniques, such as standard optical transmission/reflection13 and capacitance sensing14, scale down as the sampling area decreases, and these reduced signal levels become a concern as device features and reagent volumes reduce in size. On-chip optical refractometry15 in particular can have small sampling areas and unacceptably low sensitivity. To facilitate real-time sensing, one must improve the capabilities for localized sampling in digital microfluidic architectures.

Digital microfluidic architectures are presented in this work for enhanced actuation and sensing of fluids. A digital microfluidic multiplexer16 is employed for localized actuation through a bi-layered structure containing separated linear upper electrode columns and lower electrode rows. Multiplexing is achieved in this system at all grid locations with reduced electrical inputs through a technique utilizing differential voltage biasing. Through this technique, it is shown that multiple microdrops can be present on the device and remain stationary when only one microdrop is actuated, overcoming the multi-microdrop interference effect that has been the focus of research in path planning optimization and cross-referenced routing.7 Enhanced sensing is achieved through an optical element integrated onto the chip that allows for folded-cavity optical refractometry in a dedicated sampling station. The folded-cavity employs an overhead microlens to act as an optical back-reflection fluid monitor that is both linear and sensitive to internal fluid refractive indices. It is shown that enhanced actuation and sensing can be achieved in future digital microfluidic structures through the presented combination of digital microfluidic multiplexing and folded-cavity optical sampling.

2. DEVICE DESIGN 2.1 Fluid Actuation

Distributed electrode structures are employed in digital microfluidic devices to actuate internal fluids through the creation of spatial voltage distributions. This can be accomplished through the generalized electrode geometry of the 2-D square electrode grid.17,18 Complete control of microdrops can be achieved for all M × N grid locations, although addressability issues become apparent when highly-parallel devices are scaled to large values of M and N. For actuation, all M × N electrical inputs require an address line to be routed between electrode gaps to the internal electrodes, which become unwieldy to employ without shorting, overlapping, or crossing.

To ease the 2-D digital microfluidic constraints of addressability and enhance actuation, a cross-reference structure was introduced by Xiao et al.19 This structure uses linear electrode arrays, which are both orthogonal and separated, thus inducing 2-D horizontal and vertical movement of microdrops between the upper and lower electrodes. These linear electrodes are able to act concurrently as addressing lines and actuation electrodes, allowing for the technique to be scaled up for use with a large M × N grid. This results in M + N inputs being required instead of the usual M × N inputs. However, by reducing the number of inputs, multi-microdrop interference can occur as the applied upper and lower electrode voltages can inadvertently actuate neighbouring microdrops that share one of the same activated electrodes. To reduce this interference effect, routing,20 path planning,10 and graph theory scheduling21

have all been proposed.

An extension of the cross-referenced layout is introduced in this work with a differential AC voltage biasing technique that allows complete M × N addressability with only M + N inputs, without the interference effect of neighbouring microdrops. Voltage polarities (positive and negative) are applied to lower and upper electrodes via phase-shifts, creating a trinary input state with differential voltage amplitudes between upper and lower electrodes on the 2-D grid. This results in grounded electrodes along inactivated grid locations, a differential voltage equal to the amplitude of the applied voltage on grid locations where only one of the upper and lower electrodes is activated, and a differential voltage at twice the amplitude of the applied voltage on grid locations where both the upper and lower electrodes are activated. The known threshold voltage phenomenon22 can be exploited with this trinary state of differential voltages to initiate motion of microdrops without initiating motion in neighbouring microdrops. The voltage applied to the system must be selected to be greater than one-half the experimentally determined threshold voltage and less than the threshold voltage to allow for the isolated doubled-voltage onset of motion.

Voltage-transformed AC waveforms can be utilized to implement this threshold-based differential actuation scheme. Lower voltages are needed to actuate microdrops with AC waveforms. Furthermore, the insulating layers of the device provide low current draw, and low primary input voltages, Vin(0°), can then be used with a voltage step-up transformer. Out-of-phase AC voltage waveforms are also formed for this technique from a centre-tap transformer. The resulting digital microfluidic multiplexer with AC voltage actuation is shown in Fig. 2. The secondary side of the transformer produces positive-polarity waveforms (defined by V0(0°)) which are guided to the chosen lower electrode rows with the i-phase-switch. The secondary side of the transformer also produces negative-polarity waveforms (defined by V0(180°)) which are guided to the chosen upper electrode rows with the j-phase-switch. The inequality Vth/2 < V0 < Vth describes the connection between the amplitude of the applied voltage, V0, and the threshold voltage, Vth, where an applied voltage value within this range causes microdrop motion to be initiated at the desired crosspoint.

Figure 1. Orthogonal grids of lower and upper electrodes of the digital multiplexer are shown. The electrodes are driven by a centre-tap transformer with two out-of-phase secondary outputs, V0(0°) and V0(180°), and one single primary input, Vin(0°). The modified fluid surface tension for Vth = 0 is shown in the top figure inset, which allows for motion of microdrops along all of the activated row and column electrodes. The modified fluid surface tension for Vth ≠ 0 is shown in the bottom figure inset, which allows for motion of a microdrop at a single activated electrode crosspoint.

The microdrop motion can be quantified with activation of the ith lower row electrode and jth upper column electrode, the differential voltage at the (i, j) crosspoint, Vij, is linked to the fluid surface tension, Δγ, through23

(1)

where c is the capacitance per unit area. When the differential voltage Vij is lower than the threshold voltage Vth, microdrop motion is not induced due to the insufficient electric field. When the differential voltage Vij is greater than the threshold voltage Vth, a modified surface tension, Δγ, is created from the local electric field, resulting in microdrop motion. The localization of this actuation is made apparent in Fig. 1 in the top and bottom insets, which show a system with negligible threshold voltage and a finite threshold voltage respectively. These two insets focus on the actuated crosspoint shown in yellow in Fig. 1 at i = 12, j = 12. In the top inset, motion can exist along all actuated electrodes whereas in the bottom inset, the possibility for motion exists only at the electrode crosspoint, resulting in localized actuation. This multiplexing technique and resulting enhanced localization can greatly augment the actuation process on scalable structures and will be used for 2-D microdrop control within this work.

⎪⎩

⎪⎨⎧

≤

>−=Δ

thij

thijthij

VVVVVV

c,0

,21 22

γ

2.2 Fluid Sensing

Sensing for on-chip devices is often carried out through either optical beam interrogation24 or capacitance sampling.25 Low-frequency dielectric properties are probed through capacitance sampling through a relationship between the dielectric constant of the sampled fluid and the capacitance measured. An unfortunate drawback of this technique is that the capacitance scales down with the sampling area (reducing the scale of droplets from millimeter-scale droplets to 100 micron-scale droplets, the capacitance can reduce from 75 fF to 15 fF) 26 thus diminishing the measurement potential on smaller device dimensions. Similar to the capacitance challenge, signal powers from standard optical beam interrogation (where Fresnel reflection from the different refractive indices is utilized as a probe) diminish as the cross-sectional area is reduced. These reflected optical powers become unacceptably low with small device dimensions.

A folded-cavity sensor for refractive indices is introduced to alleviate the challenges of localized sensing. Linearity is maintained between the internal fluid refractive index and the intensity of the back-reflected optical beam for purposes of refractometry. The optical signal strength is not sacrificed at the point of sampling in this implementation when smaller device dimensions are used. Ray-tracing model results are shown on a side-view schematic of the optical sensor in Figs. 2(a) and (b). The folded-cavity optical sensor makes use of its bi-layered form and is implemented on a dedicated sampling station located adjacent to the digital microfluidic multiplexer. The incident collimated beams must converge, reflect off the lower sampling station surface, and focus within the upper glass plate image plane. This requires the microlens to be designed according to the dimensions of the multiplexer. A sensitive relationship between the refractive index of the central fluid layer and the return beam’s divergence will be used here for heightened signal intensity localized refractometry measurements.

Figure 2. Theoretical ray-tracing is shown for the folded-cavity optical sensing system is shown for the (a) air layer with nf = 1.00 and (b) filler fluid with nf = 1.55 between the two plates with (c) the corresponding simulated images from the overhead imaging camera at the detector. The required microlens design is implemented through (d) electro-dispensed fabrication of the microlens.

A back-reflected beam is shown in Fig. 2(a) with a central air layer nf = 1.00, where the focus exists within the upper glass plate, resulting in low observed optical intensities. Ray tracing results with a central fluid layer refractive index of nf = 1.55, shown in Fig 2(b), have the focus of the back-reflected beam located within the microlens, resulting in a high-intensity point focus. A noticeable contrast can be observed between the optical intensities for these internal fluids when given the appropriate focal conditions for the overhead imaging camera and a focal plane in the microlens, as shown in Fig 2(c). This system ultimately allows for the internal fluid state to be probed by the overhead camera with an improved dynamic range and heightened optical intensities.

As the microlens must accommodate the bi-layered digital microfluidic multiplexer dimensions while characterizing internal fluid refractive indices, the design of this element is critical to the fluid sensing capabilities. Electro-dispensing,27 an adaptable in-situ polymer deposition process, is employed for the fabrication of the microlens. The microlens is formed above the sampling station through this electro-dispensing technique, shown in Fig. 2(d), by dispensing a UV-curable polymer through a metallic voltage-activated dispensing tip. The profile of the dispensed droplet dictates the focusing response and is defined structurally by the nominal surface tensions of the polymer and the localized electric field that is formed between the lower grounded copper plate and the metallic dispensing tip voltage, Vdisp. The Lippmann-Young equation shows the electrocapillary response of the microlens,28

, (2)

where c is again the capacitance per unit area. The solid-filler, liquid-filler and solid-liquid surface tensions are γsf, γlf and γsl respectively, and the initial conditions, where Vdisp = 0, define the nominal contact angle of θ0. The user-controlled modified contact angle is defined as θ(Vdisp), which occurs for voltage-activated conditions where Vdisp ≠ 0. The microlens is fabricated utilizing the electro-dispensing process with particular focusing conditions dictated by Equation 2. With an appropriate substrate and surrounding filler, the microlens is given a high nominal contact angle, θ0, which is then lowered by an applied voltage on the electro-dispensing tip to tune the lens for the desired focusing conditions.

As the contact angle of the microlens is lowered, the focal length of the microlens becomes longer,29 and the nominal focus condition may be tailored based upon this dependency to create a focus within the upper glass plate when air exists in the central gap. The microlens is UV-cured when the desired form is achieved. As the fluid refractive index in the central gap is increased, the focal length will increase and drive the point of focus up towards the microlens, creating a heightened beam intensity imaged by the overhead camera.

3. DEVICE RESULTS 3.1 Fluid Actuation

The actuation structure consists of two silica plates containing copper electrodes with a thickness of 50 nm. Features are patterned via photolithography on the silica plates with an OAI mask aligner to produce a 16 linear electrode digital microfluidic multiplexer with electrode width w = 500 µm and centre-to-centre pitch p = 600 µm. Polydimethylsiloxane (PDMS) and Teflon AF are spin-coated onto each plate. Alignment of the electrode plate occurs after high-temperature curing, to produce the orthogonal form of the structure shown in Fig. 1 with a plate separation of d = 650 µm. Motion is tracked in the digital microfluidic multiplexer by an overhead high-resolution camera (LEICA APOZ6).

Figure 3. The lower row electrodes for the digital microfluidic multiplexer chip with the ith lower row copper electrodes. The chip was fabricated using the OAI mask aligner for photolithography and features copper electrodes with a thickness of 50 nm, width of w = 500 µm, and centre-to-centre pitch p = 600 µm are featured in the design above.

cos! Vdisp( ) = cos!0 +cVdisp

2

2" lf=" sf !" sl" lf

+cVdisp

2

2" lf

The voltage biasing technique is applied to the multiplexer as two opposite-phase AC waveforms at 470 Hz. A centre-tap transformer (Hammond 117E4) is utilized to generate the waveforms with a voltage gain of 2V0/Vin = 75 between the secondary terminals and the input. The i-phase-switch directs the positive-polarity secondary-side waveform V0(0°) to the ith lower electrode row and the j-phase-switch directs the negative-polarity secondary-side waveform V0(180°) to the jth upper electrode row.

Initial tests are carried out using a digital microfluidic multiplexer with a high threshold voltage, Vth = 620 Vrms, and thick (10 µm) layer of PDMS. Above-threshold motion conditions are created when the applied voltage V0 follows the 2-D multiplexing inequality Vth/2 < V0 < Vth. For the transformer being used, the input voltage of Vin follows Vth/75 < Vin < 2Vth/75. For this particular threshold voltage, an input is selected to be Vin = 8.3 Vrms, resulting in an applied voltage of V0 = 310 Vrms. Actuation of the digital microfluidic multiplexer is shown in Figs. 4(a) and (b), where microdrop 1 is initially centred at i = 6, j = 5 and microdrop 2 is centred at i = 6, j = 10. This creates a configuration where standard actuation utilizing cross-referencing would cause multi-microdrop interference along the row i = 6. In this case, however, multi-microdrop interference will not occur due to the threshold voltage requirement. Instead, multiplexing with V0(0°) on i = 8 and 9 and V0(180°) on j = 11 and 12 creates an enhanced electric field at the desired crosspoint causing preferential actuation of microdrop 2 to i = 8.5, j = 11.5, when microdrop 1 stays stationary. With only 16 + 16 = 32 inputs, complete 2-D actuation is achieved across the grid of 16 × 16 = 256 cross-points in this multiplexed structure.

Figure 4. The digital microfluidic multiplexer structure is shown with a high threshold voltage, Vth = 620 Vrms, and microdrop motion is demonstrated by microdrop 2 in the (a) initial and (b) final states, where motion does not affect microdrop 1. Activated electrodes are shown in black with the resulting activated crosspoint locations highlighted.

To increase the practicality of the device, the structure was redesigned to operate with a thin layer of PDMS and a resulting reduced threshold voltage of Vth = 48 Vrms. The multiplexer input inequality in this case is Vth/75 < Vin < 2Vth/75, requiring a voltage to be selected in the range of 0.64 Vrms < Vin < 1.28 Vrms. Actuation of microdrops in this configuration is shown in Fig. 5. The starting position for this structure has microdrop 1 centred at i = 7.5, j = 8 and a microdrop 2 located at i = 4.5, j = 12.5.

To test the digital microfluidic multiplexer, a simple algorithm for motion is tested in two separate steps of voltage activation with the results shown in Fig. 5(b). The first step is comprised of activating row i = 11 and columns j = 12 and 13 to draw microdrop 1 to the i = 10.5, j = 12.5 crosspoint. The second step consists of activating row i = 2 and columns j = 12 and 13 to pull microdrop 2 to the i = 2.5, j = 12.5 crosspoint.

A more complex algorithm, used for microdrop mixing, is shown in Fig. 5(c). The two microdrops are drawn together by activating two different crosspoint locations. Microdrop 1 is drawn to the i = 6.5, j = 11.5 crosspoint and microdrop 2

pulled to the i = 2.5, j = 11.5 crosspoint. The previous locations in each figure are shown as dashed circles. To pull the microdrops together, the voltage is then applied to rows i = 5 and 6 and columns j = 11 and 12. The end result is shown in Fig. 5(c). Motion algorithms can be programmed using a variety of activation sequences for complete 2-D multiplexed actuation.

(a) (b)

(c)

Figure 5. The digital microfluidic multiplexer structure with the adjacent sampling station is shown with a low threshold voltage, Vth = 48 Vrms, and microdrop motion is demonstrated by (a) the initial state, (b) a simple algorithm for motion, and (c) a more complex algorithm for mixing. Previous microdrop locations are shown with dotted white circles with activated electrodes in black and activated crosspoint locations highlighted.

3.2 Fluid Sensing

A folded-cavity sensor for refractive indices is incorporated as a sampling station adjacent to the multiplexer to enable localized optical sensing of fluid refractive indices between the upper and lower multiplexer plates. A collimated white light LED illuminates the structure, with a beamsplitter and overhead camera sampling the returned optical beam. The sampling station, shown in the bottom left of the Fig. 5 photographs, allows for images to be captured and the optical intensities located at the centre of the microlens to be recorded along the optical axis (OA) for various fluids.

The afore-mentioned electro-dispensing technique is utilized to fabricate the microlens on the device in order to create a well-controlled focus in the upper glass plate. The UV-curable polymer microlens, fabricated from Norland Optical Adhesive (NOA) 68, is electro-dispensed onto a PTFE-coated glass substrate. A range of dispensing tip voltages, Vdisp = 300 VDC to 1500 VDC, is used to fine-tune the microlens focus, and a final microlens with a diameter of 1.8 mm and contact angle of 50° is selected for UV-curing and device testing.

The device operation is theoretically validated by analyzing the optical response of the multiplexer and microlens with an optical ray-tracing model. Back-reflected beams are quantified at the microlens OA centre as the central fluid refractive index is changed. Theoretical model results, shown as solid squares in Fig. 6, are plotted for refractive indices ranging from nf = 1.30 to nf = 1.55. The desired linear trend is observed between the normalized intensity response and the fluid refractive indices. A normalized and unitless OA beam intensity expression, IOA(nf) = 0.49 nf + 0.15 with an R2 value of 0.996, is displayed for the linear best-fit curve.

Figure 6. Normalized OA beam intensities for ray-tracing data points (solid squares) and experimental data points (solid circles) are shown with a linear fit solid line for the ray-tracing model results. Beam patterns are shown from experimental results in the figure insets (with the OA sampling point indicated) for varying fluids and refractive indices: air (nf = 1.00), water (nf = 1.33), ethanol (nf = 1.36), and silicone oil (nf = 1.52).

The theoretical results are validated experimentally by testing the fabricated chips with fluids having known refractive indices. The fluids in the device are manipulated then directed towards the optical sampling station, where the refractive indices are then probed. The resulting optical intensity for multiple fluids, water (nf = 1.33), ethanol (nf = 1.36), and silicone oil (nf = 1.52), from the microlens OA centre is recorded and then compared to the air case (nf = 1.00). The normalized beam intensities measured in this experiment are displayed in Fig. 6 with solid circles. Strong agreement is shown between the predicted linear response and the range of fluids from nf = 1.30 to nf = 1.55. Optical beam profiles corresponding to these refractive indices are shown as insets. The optical probing in this system accurately characterizes various internal fluid refractive indices in the chip.

4. CONCLUSIONS Localized-on-chip operation of enhanced microdrop actuation and sensing applications are demonstrated in this work. The presented architecture of the digital microfluidic multiplexer allowed for decreased input complexity while providing complete 2-D microdrop motion in the system. An optical sampling station was fabricated for folded-cavity arrangement and demonstrated for highly-sensitive and local sampling of internal fluid refractive indices. Issues of scalability for digital microfluidic devices were addressed in the demonstrated work for actuation and sensing. Resulting systems can be applied for integrated technologies of future lab-on-a-chip devices.

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