Thousand-Fold Volumetric Concentration of Live Cells with a

Article pubs.acs.org/ac

Thousand-Fold Volumetric Concentration of Live Cells with a Recirculating Acoustofluidic Device Ola Jakobsson,† Seung Soo Oh,‡ Maria Antfolk,† Michael Eisenstein,‡ Thomas Laurell,† and H. Tom Soh*,‡ †

Department of Biomedical Engineering, Lund University, SE-221 00 Lund, Sweden Materials Department, Department of Chemical Engineering, Department of Mechanical Engineering, University of California, Santa Barbara, California 93106, United States

‡

S Supporting Information *

ABSTRACT: The ability to concentrate cells from dilute samples into smaller volumes is an essential process step for most biological assays. Volumetric concentration is typically achieved via centrifugation, but this technique is not well suited for handling small number of cells, especially outside of the laboratory setting. In this work, we describe a novel device that combines acoustofluidics with a recirculating architecture to achieve >1000-fold enrichment of cells in a label-free manner, at high volumetric throughput (>500 μL min−1) and with high recovery (>98.7%). We demonstrate that our device can be used with a wide variety of different cell types and show that this concentration strategy does not affect cell viability. Importantly, our device could be readily adopted to serve as a “sample preparation” module that can be integrated with other microfluidic devices to allow analysis of dilute cellular samples in large volumes.

T

mittent operation strategies, a variety of physical forces, such as hydrodynamic, electric, magnetic, and acoustic, have been utilized to trap cells within microfluidic devices.3−8 Generally, this approach offers the potential for high enrichment factors (EF) since the number of trapped cells increases continuously as a function of time (i.e., processed volume). However, this approach typically suffers from a capacity limit and a low throughput (except for ref 7) since strategies employing externally applied forces must overcome Stokes’ drag force at elevated flow rates.9 In contrast, continuous enrichment strategies do not suffer from capacity limits, as cells are constantly processed through the device. Unfortunately, it is difficult to achieve high EF with this strategy because the EF essentially governed by the ratio of the output flow rate (of the enriched fraction) to the sample input flow rate. Coakley10 et al. utilized a combination of acoustic forces and sedimentation effects to achieve 5-fold enrichment of yeast cells at 4.8 mL min−1. Seki et al.11 achieved a higher EF, showing that hydrodynamic forces could be utilized to achieve 80-fold enrichment of polymer particles, although at a significantly lower throughput (0.5 μL min−1). Recently, Nordin et al. demonstrated a device that can achieve a 200-fold EF while maintaining a throughput of 200 μL min−112 by integrating

he capability to enrich and concentrate cells and particles from dilute samples into smaller volumes is a critical step for many biological and biochemical assays. Currently, centrifugation is the most widely used technique to achieve volumetric concentration, and it is difficult to imagine the field of biology and medicine without this important method. Unfortunately, centrifugation inherently suffers from a number of disadvantages. First, it is typically a batch process performed in laboratory settings, which is cumbersome to perform at the point-of-care. Second, because of the mechanical stress it exerts on the cells, centrifugation can affect cell viability1 and phenotype.2 Finally, large reductions in volume cannot be achieved in a single step. For example, concentrating cells from a 100 mL volume into a microliter-scale volume would require multiple centrifugation steps that must be performed in series. This is especially problematic for handling small number of cells from dilute samples, because each transfer step is prone to sample loss and the pellet may not be readily visible. Over the past decade, a number of innovative approaches have been explored to achieve volumetric concentration without the use of centrifugation. Significant progress has been made in this regard by using microfluidic technology. Broadly, these experimental strategies can be categorized as being based on either intermittent or continuous operation. In both cases, the important performance metrics are enrichment factor (i.e., final concentration versus initial concentration), volumetric throughput, and maximum capacity. For inter© 2015 American Chemical Society

Received: May 25, 2015 Accepted: July 14, 2015 Published: July 30, 2015 8497

DOI: 10.1021/acs.analchem.5b01944 Anal. Chem. 2015, 87, 8497−8502

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etched in ⟨100⟩ silicon using KOH (Sigma-Aldrich, St. Louis, MO) to a depth of 150 μm and a width of 375 μm. A 1.1 mmthick borosilicate glass lid (Emmaboda Glasteknik, Emmaboda, Sweden) was anodically bonded to the silicon to seal the channel. To provide fluidic access to the chip, 7 mm long silicone tubes were attached to the inlets and outlets using silicone glue (Elastosil A07, Wacker Elastics). Acoustic Actuation and Monitoring. A piezoelectric transducer with screen-printed electrodes (PZ26, Ferroperm Piezoceramics, Kvistgaard, Denmark) measuring 25 mm × 5 mm × 1 mm (L/W/H) was glued to the chip with cyanoacrylate glue (Loctite Super Glue, Henkel Norden AB, Stockholm, Sweden). The piezoelectric transducer was actuated by a ∼2 MHz ac sinusoidal signal from a function generator (33120A, Hewlett-Packard, Palo Alto, CA), amplified by an inhouse built amplifier based on a LT1210 Op-amp, (Linear Technologies, Milpitas, CA). The amplitude of the signal was monitored by an oscilloscope (54622A, Agilent Technologies, Palo Alto, CA). A ∼10 V peak−peak sinusodial signal was sufficient to focus cells and particles within the device, which yielded a temperature increase of ∼3 °C above room temperature, measured by a PT100 sensor (Heraeus Sensor Technology, Manau, Germany) glued to the device. In order to find the frequency yielding the best acoustic focusing effect, the frequency was stepped between 1.7 and 2.1 MHz in 10 kHz steps while observing the effect in the channel. Tuning of the actuation frequency was performed when the suspension medium was changed, such as changing from DI water (microbead experiments) to a cell medium (cell experiments) but was not required when the cell type was changed. Monitoring and imaging of the microfluidic channel was done using a camera (Orca ER, Hamamatsu, Japan) mounted on a microscope (DM4000 B, Leica Microsystems CMS GmbH, Wetzlar, Germany). Sample Preparations. MCF7 and DU145 cell lines and Dulbecco’s Modified Eagle’s Medium (DMEM) was purchased from the American Type Culture Collection (ATCC, Manassas, VA), and fetal bovine serum (FBS), penicillin-streptomycin, Trypsin/EDTA solution, and LIVE/DEAD Viability/Cytotoxicity Kit for mammalian cells were purchased from Life Technologies (Carlsbad, CA). Microparticle solutions were prepared by diluting 8.3-μm fluorescent polystyrene beads (FluoroMax, Thermo-Fisher GMBH, Germany) in DI water containing 0.02% Tween20 (Sigma-Aldrich). MCF7 tumorigenic human breast epithelial cells were cultured as recommended by the ATCC in DMEM with 10% FBS and 1% penicillin/streptomycin and grown in a humidified incubator maintained at 37 °C in 5% CO2. Immediately before the experiments, adherent cells were harvested with a 0.5% Trypsin-EDTA solution and subsequently diluted in 1× PBS (Fisher Scientific, Waltham, MA) containing 0.2% BSA (Sigma-Aldrich). Human whole blood was collected from healthy donors having informed consent. Whole blood was diluted in FACS Buffer, (1× PBS, 10 mM EDTA, 1% BSA). Prostate cancer cells (DU145) were cultured and harvested like the MCF7 cells but diluted in FACS buffer prior to experiments. Fluidic Control. A syringe pump and syringe in an upright position (Ph.D. 2000, Harvard Apparatus, Holliston, MA) was used to inject the dilute sample into Inlet A on the microfluidic chip. A peristaltic pump (Reglo ICC, Ismatec, Glattbrugg, Switzerland) connected with 0.53 mm inner-diameter Tygon tubing (Ismatec) was used to pump and thus recirculate the

multiple stages of enrichment in a single microfluidic device. However, there are many applications that require higher enrichment at a higher throughput, and thus there is an unmet need for novel technologies that achieve superior performance for concentrating cells from dilute samples.7,13,14 In this work, we describe a recirculating acoustofluidic device that combines the advantages of intermittent- and continuousflow architectures to achieve, for the first time, over 1000-fold enrichment of cells at a high throughput of (>500 μL min−1). The foundation of our acoustofluidic device is a recirculation architecture that continuously increases the cell concentration over time (Figure 1A). To the best of our knowledge, such a

Figure 1. (A) Illustration of our acoustofluidic device. The dilute sample is injected into the microfluidic chip via Inlet A, while a peristaltic pump drives a recirculatory flow between Inlet B and Outlet B. Cells are focused toward the center of the channel by the acoustic radiation force and eluted through Outlet B, while the remaining cellfree solution is eluted as waste via Outlet A. This captures the cells in a recirculation loop, gradually increasing their concentration over time. (B) Photograph of the microfluidic chip. (C) Schematic of constant recirculation volume (CRV) mode. Both the syringe pump and the peristaltic pump is active. As the flow into Inlet B is matched to the flow from Outlet B, the volume stays constant in the reservoir. The concentration linearly increases in the reservoir over time for as long as fresh sample is fed into the system via the syringe (Inlet A). (D) Schematic of the reduction of recirculation volume (RRV) mode. The syringe pump is deactivated, and the peristaltic pump is active. In this mode, only 50% of the volume pumped from the reservoir is recirculated, such that the reservoir volume constantly decreases while the cell concentration increases exponentially over time.

recirculating architecture has not previously been applied to acoustofluidics. This architecture essentially has no capacity limitations because the volume of the recirculating stream is adjustable. In addition, since the acoustic force is applied perpendicular to the direction of fluid flow, Stoke’s drag does not limit the throughput as in conventional trapping systems. Finally, our system achieves exceptional enrichment because the concentration of cells continuously increases as a function of time based on a close to 100% recovery of cells in the acoustophoretic focusing step. Using our recirculating acoustofluidic device, we report over 1000-fold enrichment of mammalian cells without affecting their viability.

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MATERIALS AND METHODS Device Fabrication. The microfluidic chip was fabricated using standard photolithography and wet etching protocols, as described in the literature.15 The microfluidic channel was wet 8498

DOI: 10.1021/acs.analchem.5b01944 Anal. Chem. 2015, 87, 8497−8502

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viable cells and nonviable cells were stained by calcein acetoxymethyl ester and ethidium momodimer-1. Viability was subsequently analyzed with a microplate reader (Infinite M1000, Tecan, San Jose, CA). All experiments were performed in triplicate.

enriched sample from the bottom of the reservoir into Inlet B. The flow from the two outlets was controlled by the hydrodynamic resistance from two pieces of 35 cm-long 300μm (i.d.) tubing (Sigma-Aldrich). As a result of the approximately equal hydrodynamic resistance, the flow was split ∼50-50 between the center and side outlets. An important consideration for our device design was to optimize the flowsplit ratio between the two outlets to achieve minimal cell loss while maximizing volume throughput (see section SI.2 in the Supporting Information). The flow from Outlet A was collected in a large waste tube and the enriched sample from Outlet B was collected in the reservoir, as illustrated in Figure 2. In both CRV and RRV modes, the peristaltic pump flow-rate was 1 mL min−1 for MCF7 cells and microbeads and 0.5 mL min−1 for RBCs and DU145. For CRV, the syringe pump flowrate was set to match the flow from the peristaltic pump, resulting in a total flow-rate in the microchannel equal to double the flow from the syringe pump. The sample throughput in CRV mode is equal to the syringe pump flowrate. The sample throughput in RRV mode equals 50% of the peristaltic pump flow rate. Depending on the sample, reducing 1 mL of sample to approximately 100 μL (10-fold enrichment) of enriched sample took between 2 and 4 min in RRV mode. Sample Concentration Measurements. A flow cytometer (FACS Verse, BD Biosciences, San Jose, CA) equipped with a flow meter was used to analyze the concentration of the microbead and MCF7 cell samples. The sample was first run through the cytometer for 30 s in order to stabilize the flow before measurements were taken. A minimum of 240 μL of sample or 10 000 events was analyzed for each sample. For samples with a volume less than 240 μL (as was typically the case for the enriched samples), 50 μL of the sample was diluted in 1 mL of DI water or PBS. Fluorescent beads were identified and gated in the software (BD FACSuite) using a dot plot, FSC-A vs FL1-A. Cells were gated and identified in a SSC-A vs FSC-A dot plot. For RBC and DU145 cell samples, a BD Accuri A6 (BD Biosciences) was used to determine the cell concentration, with cell counting performed according to the manufacturer’s manual. Standard series with known concentrations of the specific cells or beads used for the experiments were run through both instruments to verify the measurement method and accuracy. The enrichment performance was calculated using enrichment =

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RESULTS AND DISCUSSION Device Design and Modes of Operation. Our device uses a continuous-flow, acoustofluidic design to create standing

Figure 2. (A) Bright field image of the flow splitter region of the acoustofluidic chip. Parts B, C, and D show the same region imaged by fluorescence microscopy at different times during a 1 h run with our device operating in CRV mode. The outline of the microchannel is highlighted in red. A faint band of acoustically focused fluorescent 8.3μm beads can be seen exiting through the central outlet, gradually increasing in concentration (and hence fluorescence intensity) over time.

Figure 3. (A) Average enrichment factors obtained in CRV mode from four different starting volumes. The input sample was processed at 1 mL min−1 and the volume of the enriched fraction was kept at 1 mL. The enrichment factor is a linear function of processed volume. (B) Average enrichment factors obtained in RRV mode after concentrating a single starting volume (1 mL) to four different final volumes. All experiments were started with 1 mL of sample in the reservoir. Error bars in both panels were obtained from three separate experiments.

enriched sample concentration initial sample concentration

Recovery was estimated by measuring the concentration of cells eluted through the waste outlet and calculated according to recovery = 1 −

Table 1. Enrichment Data from Sequential Modea

waste outlet concentration initial sample concentration

This approximation neglects losses of cells derived from “sticking” in the system and thus overestimates the recovery. Visual inspection of the microfluidic channel after an experiment indicated that very few cells were stuck in the channel. Cell Viability Measurements. Cell viability was measured with the LIVE/DEAD Viability/Cytotoxicity Kit for mammalian cells (Life Technologies), and the measurement was performed according to the manufacturer’s protocol. Each cell sample was transferred into a black 96-well microplate, and

starting volume (mL)

avg starting concentration beads (μL−1)

avg enriched fraction concentration beads (μL−1)

avg enrichment factor

avg recovery

60 120

0.91 0.84

644 1289.75

720 ± 176 1534 ± 69

95 ± 3% 96 ± 2%

a Average values and standard deviation are derived from three separate experiments.

waves in the microfluidic channel, generating an acoustic force (see section SI.1 in the Supporting Information) that laterally displaces cells to the center of the channel (Figure 1A, B). This device design has been previously used by our group and 8499

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In practice, the EF is limited by the final volume that can be eluted without loss of cells. The final mode of operation is a “sequential mode”, wherein the CRV and RRV modes are executed consecutively to achieve the highest enrichment. In other words, all of the cells are first captured in the reservoir using the CRV mode as described above. Then, the total volume in the collection reservoir is reduced using the RRV mode. In this way, exceptionally high EF can be obtained, because the EFs obtained with the CRV and RRV modes can under ideal conditions be multiplied as shown below: EFsequential = EFCRV × EFRRV

Figure 4. Enrichment of red blood cells (RBCs) at a flow-rate of 1 mL min−1. The end of the microfluidic channel (including part of the flowsplitter) is imaged without (A) and with (B) acoustic actuation. Without acoustic actuation, the RBCs remain dilute and visually indistinct. With acoustic actuation, however, the enriched cells are highly concentrated in the center of the channel to form a focused red stream with clear borders.

Characterization of the CRV and RRV Modes Using Synthetic Particles. We initially characterized the performance of our device in CRV mode, using polystyrene beads (diameter = 8.3 μm) suspended in DI at very low concentrations (∼1 bead/μL). The sample was loaded into a syringe pump and injected into the device at 1 mL min−1; the peristaltic pump was operated at a flow-rate equal to the syringe pump, making the total flow-rate in the device 2 mL min−1. The fluid volume within the recirculation loop through the reservoir was kept constant at ∼1 mL. We could visualize the enrichment of the beads by microscopically observing their fluorescence intensity. Qualitatively, the fluorescence intensity increased approximately linearly due to the linear increase in bead concentration as a function of time (Figure 2). We subsequently determined the EF after performing experiments with four different initial volumes (10, 30, 65, and 100 mL) and measuring the concentration of beads before and after enrichment using flow cytometry (see Materials and Methods). For all initial sample volumes, the enrichment performance of our system in CRV mode matched very well with theoretically expected values (Figure 3A). Next, we characterized the performance of our device in RRV mode. Briefly, 1 mL of sample containing polymer beads (∼15 bead/μL) was loaded in the reservoir and then reduced to either 750, 500, 200, or ∼100 μL. The peristaltic pump was operated at 1 mL min−1, and the volume in the reservoir was reduced at a rate of 500 μL min−1. As shown in Figure 3B, the experimental and theoretical values of EF matched well (i.e., exponential enrichment). The highest EF of 10 was obtained when 1 mL of the initial sample was reduced to 100 μL. We note that the absolute minimum collection volume with our system is the internal volume of the fluidic system (the volume of the microfluidic chip and tubing), which is ∼50 μL. Because of the larger experimental error in measuring smaller collection volumes, the standard deviations were larger for higher EF values. Characterization of Sequential Mode Using Synthetic Particles and Mammalian Cells. The sequential mode entails the tandem operation of CRV and RRV modes to achieve the highest EF. As a demonstration, we first processed a

others.15,16 Our recirculating acoustofluidic device can be operated in three different modes. The first is “constant recirculation volume” (CRV) mode, in which cells in suspension are injected into Inlet A of the chip with a syringe pump (Figure 1C). Because of the acoustic standing wave, all cells are focused to the channel center and are recovered through Outlet B into the recirculation reservoir. Importantly, a peristaltic pump transports the contents of the reservoir back into Inlet B at a rate that maintains the reservoir at a constant volume of typically 1 mL. In this way, the recirculation loop essentially serves as a “microfluidic trap” with continuously increasing cell concentration. Assuming full recovery of cells, EF can be expressed as EFCRV =

processed sample volume final volume in reservoir

The second type of operation is “reduction of recirculation volume” (RRV) mode, in which the sample is loaded into the reservoir and a peristaltic pump is used to inject the sample into Inlet B of the chip (Figure 1D). Importantly, only a fraction (typically ∼50%) of the injected volume is collected through Outlet B, which contains all the cells in the sample as a result of the acoustic focusing. The remainder of the volume, which contains no cells, is eluted through Outlet A and discarded. In this way, the volume of the reservoir is continuously reduced while the number of cells ideally remains constant, effectively making the EF increase exponentially as a function of time. Once a desired volume is reached in the reservoir, all of the cells in the chip and tubing are eluted with an air bubble and collected into the reservoir for analysis. Thus, the EF of the RRV mode can be expressed as EFRRV =

initial sample volume in reservoir final volume in reservoir

Table 2. Enrichment Data from Sequential Mode Operation from Various Human Cell Typesa cell type

processed volume (mL)

processing time (min)

avg starting sample conc (cells μL−1)

approx. collected volume (μL)

sample flow rate (mL min−1)

enrichment factor

recovery (%)

RBC DU145 MCF7

100 100 60

200 200 60

56.7 4.3 1.3

100 100 100

0.5 0.5 1

1166 ± 110 817 ± 125 519 ± 115.7

98.7 ± 0.1 90.2 ± 3.1 81.7 ± 15

a

Average values and standard deviations are derived from three separate experiments. 8500

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Figure 5. (A) Acoustofluidic concentration does not affect cell viability. Live/dead cell assays show essentially equivalent viability for untreated MCF7 cells (blue) compared to those that have passed through the peristaltic pump alone (red) or through the microfluidic chip in the absence (green) or presence (purple) of acoustic focusing. Parts B−D show micrographs of proliferating MCF7 cells after enrichment. MCF7 cells enriched ∼500-fold by our device were grown in culture medium at 37 °C. Over 3 days, we observed robust growth of adherent, polygonal cells with clear cell boundaries, confirming cell health and viability.

force (See section SI.1 in the Supporting Information) on the DU145 cells, as compared to the RBCs. Although we were able to enrich MCF7 cells at a higher throughput (1 mL min−1 vs 0.5 for the other cell types) and achieve an enrichment factor larger than 500, we also observed a lower recovery (81.7% ± 15), presumably due to the higher flow rate. The higher flow rate (and thus shorter duration of the experiment) was used to reduce the time-dependent nonspecific aggregation of the cells that confounds flow cytometric analysis.17 Our device is expected to function with similar performance for most mammalian cells,18 and over a wide range of concentrations.19 Much smaller cell types, however (such as bacteria), is subject to a lower acoustic radiation force20 and thus requires a longer exposure time in the acoustic field in order to be deflected and recovered in the recirculation loop. As the acoustic exposure time is a function of the flow rate, a lower flow rate would be required for the enrichment of smaller cell types, in order to maintain a high recovery and EF. Finally, we evaluated the effect of our device with respect to cell viability using a commercially available live/dead cell assay (see Materials and Methods). As shown in Figure 5A, cells enriched in our device showed minimal difference in viability in comparison to untreated cells, cells pumped through the peristaltic pump only, or cells processed through the chip without ultrasound. Importantly, enriched cells proved highly viable and capable of proliferating when incubated in culture medium over the course of 3 days (Figure 5B−D)

highly dilute, high-volume sample in CRV mode to capture all of the cells into the reservoir, maintaining a total volume of ∼1 mL at a sample-in flow-rate of 1 mL min−1. Subsequently the syringe pump was turned off and the device was operated in RRV mode to reduce the total volume to approximately 70− 100 μL. The concentration of the cells before and after the experiment were measured using a flow cytometer as described above. To characterize the performance of our device in the sequential mode, we used a highly dilute sample (1500. The sequential mode can also be utilized with a wide variety of live cells to achieve unprecedented enrichment at high throughput. We evaluated the performance of our device in sequential mode with diluted samples of human breast cancer cells (MCF7), red blood cells (RBCs), and prostate cancer cells (DU145). As with the polymer beads, we could readily image the enrichment of the cells using brightfield microscopy (Figure 4). We then obtained the recovery and EF for each of the three cell types using flow cytometry. The experimental conditions and data are presented in Table 2. We achieved the highest EF (1166 ± 110) and recovery (98.7% ± 0.1) with RBCs at a throughput of 0.5 mL min−1. With DU145 cells, we also achieved exceptional enrichment (817 ± 125) and recovery (90.2% ± 3.1). The difference in recovery can be explained by a slightly lower acoustic radiation

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CONCLUSION In this work, we describe a simple and powerful recirculating scheme for an acoustofluidic device that combines the advantages of intermittent- and continuous-flow architectures, 8501

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(16) Yang, A. H. J.; Soh, H. T. Anal. Chem. 2012, 84, 10756−10762. (17) Shapiro, H. M. Practical Flow Cytometry, 3rd ed.; Wiley-Liss, Inc.: New York, 1995; p xxxviii. (18) Burguillos, M. A.; Magnusson, C.; Nordin, M.; Lenshof, A.; Augustsson, P.; Hansson, M. J.; Elmer, E.; Lilja, H.; Brundin, P.; Laurell, T.; Deierborg, T. PLoS One 2013, 8, e64233. (19) Lenshof, A.; Ahmad-Tajudin, A.; Järås, K.; Swärd-Nilsson, A.M.; Åberg, L.; Marko-Varga, G.; Malm, J.; Lilja, H.; Laurell, T. Anal. Chem. 2009, 81, 6030−6037. (20) Antfolk, M.; Muller, P. B.; Augustsson, P.; Bruus, H.; Laurell, T. Lab Chip 2014, 14, 2791−2799.

demonstrating >1000-fold enrichment from starting volumes of up to 100 mL at an exceptional volumetric throughput (>0.5 mL min−1) and with a high recovery of cells. Importantly, we demonstrate that our device has negligible effect on the viability of the enriched cells as shown by both live/dead assays and proliferation in culture. We envision many possible applications of our device for stand-alone operation to replace centrifugation, especially for processing very small number of cells from large volumes. Importantly, we also believe that our device could be readily modified and integrated into a larger microfluidic system as a “front-end module” for sample preparation. Such an assembly could form the foundation for a truly integrated system to perform molecular analysis of rare cells directly from highly dilute samples at the point of care.

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ASSOCIATED CONTENT

S Supporting Information *

Introduction to the physics and governing equations to the acoustic radiation force (section SI.1) and a more detailed description on how the flow rates were chosen and optimized for the recirculating device (section SI.2). The Supporting Information is available free of charge on the ACS Publications website at DOI: 10.1021/acs.analchem.5b01944.

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AUTHOR INFORMATION

Corresponding Author

*E-mail: [email protected]. Phone: 805-893-7985. Fax: 805893-8651. Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS We are grateful for the financial support from The Garland Initiative, DARPA (Grant N66001-14-2-4055) and the Institute of Collaborative Biotechnologies through the Army Research Office (Grants W911NF-09-0001 and W911NF-10-2-0114).

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REFERENCES

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DOI: 10.1021/acs.analchem.5b01944 Anal. Chem. 2015, 87, 8497−8502