point-of-care, portable microfluidic blood analyzer...

Point-of-care, portable microfluidic blood analyzer system

Teimour Maleki1, Todd Fricke4, JT Quesenberry4, Paul W. Todd4, James F. Leary1-3 1Birck Nanotechnology Center, 2 Weldon School of Biomedical Engineering,

3Department of Basic Medical Sciences, Purdue University School of Veterinary Medicine,

Purdue University, West Lafayette, USA,

4Techshot, Inc. Greenville, Indiana, USA

ABSTRACT

Recent advances in MEMS technology have provided an opportunity to develop microfluidic devices with enormous potential for portable, point-of-care, low-cost medical diagnostic tools. Hand-held flow cytometers will soon be used in disease diagnosis and monitoring. Despite much interest in miniaturizing commercially available cytometers, they remain costly, bulky, and require expert operation. In this article, we report progress on the development of a battery-powered handheld blood analyzer that will quickly and automatically process a drop of whole human blood by real-time, on-chip magnetic separation of white blood cells (WBCs), fluorescence analysis of labeled WBC subsets, and counting a reproducible fraction of the red blood cells (RBCs) by light scattering.

The whole blood (WB) analyzer is composed of a micro-mixer, a special branching/separation system, an optical detection system, and electronic readout circuitry. A droplet of un-processed blood is mixed with the reagents, i.e. magnetic beads and fluorescent stain in the micro-mixer. Valve-less sorting is achieved by magnetic deflection of magnetic microparticle-labeled WBC. LED excitation in combination with an avalanche photodiode (APD) detection system is used for counting fluorescent WBC subsets using several colors of immune-Qdots, while counting a reproducible fraction of red blood cells (RBC) is performed using a laser light scatting measurement with a photodiode. Optimized branching/channel width is achieved using Comsol Multi-PhysicsTM simulation. To accommodate full portability, all required power supplies (40v, ±10V, and +3V) are provided via step-up voltage converters from one battery. A simple on-board lock-in amplifier is used to increase the sensitivity/resolution of the pulse counting circuitry.

1. INTRODUCTION As we1-5 and others6-8 have previously proposed, microfluidic flow cytometers have the potential to be

inexpensive closed-system, portable, point-of-care, diagnostic devices. However, many current microfluidic cytometers are �“micro�” only in the microfluidic chip whereas the rest of the apparatus is decidedly �“macro�”, usually requiring microscopes or other large ancillary optical or electronic equipment, and a separate computer9-14. In this paper we report on progress in developing a hand-held device designed to be a truly portable, battery-powered and hand-held device for point-of-care medical applications that can be operated by untrained, or minimally trained personnel. Existing commercially available flow cytometers are still relatively large, not very portable, expensive, and still require highly trained personnel to operate correctly, mainly because they are general-purpose devices that require expertise to adapt them to specific applications. We believe the expertise should be built into a smart device that already knows how to process samples and the incoming data in real-time by using embedded algorithms.

Microfluidic flow cytometers, if properly configured and truly portable in entirety, can solve many of these problems. A microfluidic cytometer can be easily made as a closed-system �“reader�” device with disposable chips reducing cost, increasing sterility, and preventing contact with hazardous agents by doing

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on-chip cell preparations with insertable reagent packs that can even be designed to talk to application-specific embedded algorithms through barcoded reagent chips (Figure 1). While this last feature is not yet implemented in our current prototype, it could be fairly easily implemented making the device software configurable to do related, but differing, tasks. The overall true portability of our device should allow for operation outside a traditional medical setting, further reducing possibilities of infection or contamination occurring from sample preparation and transport.

2. MATERIALS AND METHODS

2.1. System Design

Three-dimensional design of the microfluidic cytometer is illustrated in Figure 2. The total dimension of the box is 8×16×18 cm3. As depicted in the figure, the cytometer is composed of four major units: electronic unit including battery and function-specific electronic processing boards; fluidic handling including the microfluidic chip, static mixer, magnetic cell sorter, buffer/reagent reservoirs and actuator pump; optical handling including ultraviolet light emitting diodes (UV-LEDs), collection/focusing lenses, and optical filters; and user interface. A cross section of the optical unit along the detectors is depicted in Figure 2-b showing four optical paths and the locations of the lenses and filters. Excitation lenses collect the UV-LED light and focus it on the microfluidic channel, while the emission lenses collect the fluorescence signal and focus it on the avalanche photodiode (APDs) (aka silicon photomultipliers(SPM)) detectors. A closer view of the optical/fluidic processing unit is shown in Figure 2-c. The door that the chip gets inserted through has a key shaped opening that requires the valve to be turned to its �“in-use�” position to load it into the cytometer. Turning this handle will actuate the magnet and bring it close to the microfluidic channel, Figure, 2-d. The rest of the process is automated and begins when the user closes the door and engages the door lock.

Figure1: Conceptual schematic of the subcomponents and processes that constitute an integrated portable microcytometer reader, reagent packs, and disposable microfluidic chip. White bloods cells are separated from red blood cells by immunomagnetic, valveless sorting and separately counted down different microfluidic paths. White blood cells are further processed by immunophenotyping using three different colors of commercially available antibody-conjugated Qdots (quantum dots).5

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2.2. Disposable Cartridge

A schematic of the disposable cartridge is depicted in Figure 3. As can be seen from Figure 3-a, the disposable chip is composed of two parts: up-front fluid processing and microfluidic chip. The up-front fluid processing unit is fabricated by rapid prototyping using Stereo Lithography Apparatus (SLA) (Acu-Cast Technologies, TN, USA) while the microfluidic branching system (blue plate in the figure) is fabricated using standard molding of Polydimethylsiloxane (PDMS) against an SU-8 mold. In the SLA process, the layering resolution is 50 m, and the minimum channel width is 125 m. Detailed views of the microfluidic branching system have been previously reported1-3. The white piece is a retainer to hold the plungers in place. Buffer will be loaded into the two pistons all the way up to the valve (red passages) and reagent will be loaded in the large passage (purple) thru the valve via the septum. This will be done �“at the factory�”. The user/patient places a lancet-punctured finger at the sample inlet on the side of the cartridge (green), turns the valve 90 degrees and removes the retainer before inserting the cartridge into the cytometer. There are stops to limit the valve to only these two positions. The door that the chip gets inserted through has a key-shaped opening that requires the valve to be turned to its �“in-use�” position to load it into the cytometer. The rest of the process is automated and begins when the user closes the door and engages the door lock. Yellow passages are the drains and they are vented to atmosphere by means of a hydrophilic membrane.

A detailed view of the front end of the cartridge is shown in Figure 3-b. The rotating blue stopcock is set to the �“operate�” position so that the sample and reagent are forced, side-by-side to flow into the static mixer, which has 6 mixing stages. The driving force is the two channels of buffer flow induced by the actuator drive, which pushes two pistons at the same velocity. The cross section of the sample channel is 1/10th the cross section of the reagent channel so that, for example, 1 L of sample is combined with every 10 L of reagent as the fluids enter the mixer. Reagent and buffer are expected to be pre-loaded prior to operation while the user adds sample (a droplet of whole blood).

Figure 2: (a) schematic of the total system showing different components of the system, (b) cross sectional drawing of the optical unit at the location of the detectors, (c) combined optical fluidic processing unit, (d) bottom view of the system showing the magnetic actuator system.

(a) (b)

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A close-up view of a single stage of the static mixer is demonstrated in Figure 3-c. The mixer design is called an Interfacial Surface Generator (ISG). This design is available commercially from Ross Static Mixers (Hauppauge, NY, USA); however no commercial mixer is available in the 100 L range. We have undertaken the design of the smallest ISG ever made. As can be seen in the picture, each mixing element has four channels connecting tetrahedral chambers. Fluid enters four channels aligned on a vertical edge of the left tetrahedral chamber. Fluid is deflected by the four channels so that it emerges from the four channels aligned on the horizontal edge of the right tetrahedral chamber. This brings the fluid from the four channels together in parallel �“bands�”, so that the first mixing stage produces 8 bands of the 2 fluids; the next stage rotates the bands 90° and passes them to the next chamber in 32 bands etc. so there will be 2 x 4n bands in the nth stage. We have designed a 6-stage mixer, which results in 8,192 bands with maximum unmixed fluid width of 1000/8192 = 0.12 m, or less than 1/10th the diffusion distance of a whole cell in 1 second (keeping in mind that a whole cell is generally considered �“non-Brownian�”.

2.3. Electronic Requirements

For proper operation, the cytometer needs electronic detection circuitry, UV LED excitation, APD detectors, LCD user interface, and digital processing and control systems. These imply that there should be four main power subsystems: power for the SPMs (30V, 20 µA), power for the LEDs (3.5 V, 4×0.5 A), power for electronic readout circuitry (±10 V, 50 mA), and power for the processing unit (±3.3 V, 50 mA) analog devices to the circuitry in the system. It should be noted that the output signal is sensitive to the SPM�’s supply voltage and UV LEDs current. Hence, SPMs must be powered with a fixed voltage with low noise while LEDs must be powered with fixed current. Furthermore inherent SPM dark noise cancelation is cumbersome when working with a DC voltage; therefore, we powered the UV-LED with a 100 kHz square wave (50% duty cycle) and implemented a simple on-chip lock-in amplifier to overcome this issue. The

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Figure 3: Detailed views of the fluidic system: (a) structure of the disposable microfluidic chip contains microfluidic channels, static mixer, buffer/reagent reservoirs, loading valve, and drain channels; (b) closer top view of the upfront fluidic processing; (c) 3-D schematic of single stage of the ISG mixer.

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Qdot® 705 conjugate (Q10059, Invitrogen, NY, USA). Normalized excitation/emission of the above reagents is depicted in Figure 6. As can be seen from the graph, Qdots have a very narrow excitation band while Hoechst has a wider excitation band. It is worth mentioning that generally Hoechst is much brighter than Qdots; therefore, even though the normalized graph below shows that Hoechst emission has low intensity at 600 nm, it could be still brighter than Qdot 605 excitation. Another important finding in Figure 5 is that excitation at 365 nm would result in brighter Hoechst compared to Qdots. For optimized results, one could use lower excitation wavelength (such as 325 nm); nevertheless there is currently no bright commercial LED at such wavelengths.

2.5. Optical Components

Ultra high power UV LEDs (NCSUO33A, Nichia, Tokyo, Japan) were used for the excitation light while 1 mm SPMs with 100 m microcell size (PMMicro1100X18, SensL Technologies, Ltd, Cork, Ireland) were used for detectors. For the optical unit we used N-BK7 Plano-Convex lens (LA1116, Thorlabs, NJ, USA) as the focusing and collection filters, and custom-made filters from Omega Filters (400ALP, 750SP, 450BP25, 650BP25, and 605BP25, Omega Optical, Inc, VT). The 750SP filter was used for rejecting near infrared heat from the UV LED. .

3. EXPERIMENTAL RESULTS 3.1. Microfluidic Chip

As previously reported, the microfluidic system is a two-stage branching system1-3. An optical image of the microfluidic chip is depicted in Figure 7-a. The main channel is 300 µm wide 30 µm deep and 30 mm long. These dimensions are chosen for optimum pressure/velocity of the cells in the channel for reasonable processing time (10 minutes) and effective magnetic WBC sorting. The first branch, into which magnetically labeled WBCs are diverted, is 100 µm wide, 30 µm deep and 25 mm long. The cross sectional

Figure 6: Normalized intensity versus wavelength of the excitation/emission of the staining reagents used in the portable microfluidic cytometer. The bandpass filters used for collection of specific fluorescent colors are in italics and are approximately depicted by the shaded rectangles.

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dimensions and area of this microchannel were chosen based on Comsol MultiphysicsTM simulations so that 10% of the blood volume is diverted into this branch. The length of the channel was chosen to comprise four detections sites. The second branching (70 µm in width, 30 µm in depth, and 5 mm in length) is devoted to counting of a sample (10%) of RBCs. The RBC population in blood is so high (~109/ mL) that counting all the RBCs in 10 minutes requires ultrafast cell counting which is not simple to implement in a handheld cytometer. Furthermore, since we are counting a repeatable volumetric fraction of the sample, we back calculate the total number of the RBCs.

An issue usually associated with microfluidic system is adhesion of the blood cells to the channel�’s sidewall. To characterize this phenomenon, whole blood diluted in PBS (10% whole blood + 90% PBS) was pumped into a dry, untreated microfluidic system with flow rate of 20 µl/min for 10 minutes. Immediately after diluted blood, PBS was pumped into the microchip with a flow rate of 20 µl/min for 10 minutes. No leakage was observed during the test, and microscopic evaluation of the system afterwards did now show any blood cells in the channel, see Figures 7-b to d.

3.2. Magnetic Cell Sorting

In the previous report4, the requirement for the magnet to reside in very close proximity to the main channel for successful magnetic collection of the WBCs was demonstrated. However, this is technically difficult to achieve with a disposable PDMS microfluidic. The magnet is too expensive to be a part of the disposable cartridge. We therefore optimized the magnet location below the channel by carving a location for the magnet from the backside of the microfluidic system using laser micromachining, as discussed earlier, an actuation mechanism was developed to bring the magnet into close proximity of the channel from the backside. Figure 8 demonstrates the time-exposure images of the integrated trajectories of WBCs (WB flow rate=2 µl/min). As shown in Figure 8 (a), a cell is drawn from a position about 2/3 of the way across the channel. A cell is traveling in an S-shaped trajectory to enter the WBC branch, (b), an indicator of the success of the magnetic sorting. Magnetophoretic mobility (measured by particle tracking velocimetry) of WBCs labeled with anti-human CD45 magnetic microspheres was 1.19 x 10-11 m3/TAs,

Figure 7: (a) Photograph of the microfluidic chip before connecting to the mixer unit. Passage of the whole blood in untreated sample did not leave any residue in the channel. (b), (c), and (d) shows optical microscopy image of the microfluidic system after whole blood passage at the first branch, second branch and in the RBC channel respectively.

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A photograph of the optical processing unit and the cartridge holder is shown in Figure 9-a. Location of the magnet, door lock mechanism and magnet actuation system can be seen in this figure. A component view of the optical unit encompassing the UV-LEDs and laser photo diode at top, APDs and photo-detector at bottom and lenses/filters at the middle in an anodized aluminum package is demonstrated in Figure 9-b. As can be seen in Figure 9-c, disposable cartridge (gray plastic) will reside in a slider with rails, and door lock will rotate 90-degrees to hold the cartridge in place and bring the magnet into close proximity of the channel. The digital display registers and displays counts in all five channels and displays accumulating events in real time, Figure 9-d.

Figure 10 shows photographs of the electronic boards employed in this system. The outputs of the SPMs are fed into the detector board, Figure 10-a. This filters noise, amplifies the signal and adjusts the signal level to be processed by analog to digital converter. The data acquisition board accepts signals from the electronic read-out and passes them to the display, Figure 10-b. The �“Interposer�” circuit board consists of connectors and protective diodes so that the power and signal boards can transfer signals from the amplifier circuits to the alphanumeric display; it serves as a power bus. The power board controls the pulse pattern of the LEDs and sends the synchronous data for pulse conditioning of the detector board. Figure 10-c.

(c) (b)(a) Figure 10: Photographs of: (a) the detection board circuitry including on board lock-in-amplifier, (b) the powering board and the user interface, (c) controller board that will be attached to the bottom of the power board.

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Figure 9: Photographs of (a) the optical detector unit composed of three different stages (excitation UV LED�’s and the lenses, collection lenses and emission filters, and SPM detectors), (b) the chip holder section showing the location of the magnet, and assembled optical path, (c) microfluidic system in place, (d) The digital display registers and displays counts in all five channels and displays accumulating events in real time.

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3.4. Optical Detection System The detection of SortCalTM beads (iCyt Mission Technology, Champaign-Urbana, IL, USA) in the

microfluidic channel using larger lenses and filters (25mm diameter) was reported previously1-3. Similarly, to test the new optical detection system with smaller lenses and filters, a 0.2 mm wide 3 mm long slit was placed above the WBC channel and the excitation LED was powered with a 100 kHz square wave with 50% duty cycle. Figure 11 shows the SPM signal output before and after processing with the lock-in-amplifier. Pulsing the light is very important as it allows sampling and real-time subtraction of underlying noise as well as integration over lock-in amplifier specified time intervals. As can be seen, the use of lock-in amplifiers greatly enhances the signal-to-noise ratio.

3.5. Mixer Characterization

The mixer was coated with PEG (PEG filling was almost bubble free), and whole blood, treated with ACD as anticoagulant, and the magnetic labeling reagent (BioMag® anti-Human CD45 (BM588/10114) diluted 1:5 with PBS) was introduced to the inlets of the mixer, Figure 12. The labeling reagent flow rate was set to be 20 µl/min while the whole blood flow rate was 2 µl/min. After 10 minutes, the whole blood was replaced by PBS (flow rate=20 µL/min), and the flow rate of the reagent was set to zero. This will block the upper inlet, while PBS will be used to flush the mixed blood from the mixer. At the end of the experiment, the mixer was flushed with 10% bleach and 10% ethanol, and DI water with high flow rates to completely clean the mixer to be reused again. Fluid from the outlet of the mixer (~220 µL) was processed by passing it through a magnetic trap to separate magnetically labeled white blood cells from the red blood cells. The trap was flushed with 500 µL of PBS while the magnetic was in position. This collected material is considered a negative control. After removing the magnet, the trap was rinsed with 1 mL of PBS, and the collected liquid is considered positive control. Ideally, the negative control should not have any white blood cells and the positive control should not have any red blood cells. In fact there is a small fraction of remaining RBC in the WBC fraction but this will have no effect on the device�’s final WBC counts since the RBCs are non-nucleated and will not be counted since they will not stain with the DNA �–specific Hoechst 33342 dye. Likewise the small number of WBC that will not be sorted down the WBC channel but will appear in the RBC channel will have a negligible effect on the RBC count since they will constitute a tiny fraction (<< 1%) of the total RBC channel count.

Figure 11: Qdot labeled cell detection with SPM with WBCs flowing in the channel at 10 L/min. Use of lock-in amplifiers synchronized with the strobed UV LED excitation source greatly increases the S/N ratio.

Before Lock-in amplifier

After Lock-in amplifier

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4. CONCLUSIONS

Portability is the main focus of this microfluidic cytometer, and we have made strides in designing and

testing elements of a portable cytometer. The PDMS microchip with magnetic sorting is simplistic and performs reliably and reproducibly. The optical system is small enough using a combination of LEDs, SPMs, and 6 mm lenses. Currently a benchtop lock-in amplifier is being used but that will be replaced with a custom hybrid lock-in amplifier consisting of simple analog circuitry and a programmable microchip. Pulsing the LED light source saves power. We calculate that the device will be able to run using a lithium battery similar to those used in cellular telephones. A front-end design that uses capillary uptake to obtain a one-drop blood sample and then a simple stepper motor driving pistons to mix the sample and drive fluid through the chip is also a feature. Ultimately at 10 L/min rates, a complete run of the micro cytometer would take less than 30 minutes including on-chip sample preparation, cell staining, data acquisition and automated data analysis. When completed this micro cytometer will be one of the first truly portable, point-of-care diagnostic devices for whole blood cell analysis.

Design, development and characterization of elements of a portable microfluidic cytometer have been presented. All the components are powered from a single rechargeable lithium battery, similar to those used in portable personal electronics. A simple multi branching microfluidic system was fabricated by molding PDMS against a SU-8 master mold. On-chip magnetic sorting was successfully employed to concentrate the white blood cells after selective magnetic labeling. All the optical components (including the UV-LEDs, APD detectors, lenses, and filters) are small enough (less than 7 mm in diameter) to be integrated into a hand-held housing. A simple on board lock-in-amplifier was designed to enhance the signal-to-noise ratio by reducing the dark noise of the SPMs. Exciting the UV-LEDs with a pulse rather than DC voltage resulted in reduced power consumption and enhanced sensitivity. Blood sample processing is realized by combining a drop of blood taken into a built in capillary, and a miniaturized stepper motor is used to mix the sample and drive the fluid through a mixer and the microfluidic chip. Ultimately at the rate of 20

L/min, the cytometer needs only ~30 minutes to complete the processing of blood. A color LCD displays counts proportional to total white blood cells, total red blood cells and counts of three WBC subsets such as those carrying CD8, CD4 and CD11b biomarkers.

ACKNOWLEDGEMENTS

The authors thank the Birck Nanotechnology Center and Bindley Bioscience Center staffs for their help

and support. We specifically appreciate the guidance of Ms. Lisa Reece, Christy Cooper, and Trisha

Magnetic reagents: flow rate=20 µl/min

Whole blood: flow rate=0 µl/min

Figure 12: Whole blood labeling with BioMag® anti-Human CD45 (diluted with PBS 1:5) using the mixer. (a) Mixer was coated with PEG in advance and was filled with reagent before introducing the blood into the mixer. (b) Blood is also introduced in the second inlet to be mixed with the reagent that continues to be introduced in the first inlet.

Blood traces in the mixer

Mixed blood

Magnetic reagent

(a) (b)

Whole blood: flow rate=2 µl/min

Magnetic reagents: flow rate=20 µl/min

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Eustaquio for their help in blood processing. This work was funded by NASA contract NNX10CB06C Techshot, Inc. and a subcontract to Purdue University issued by Techshot, Inc.

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