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Biomicrofluidics. 2013 Mar; seven(ii): 024103.

A negative-pressure-driven microfluidic chip for the rapid detection of a float cancer biomarker in urine using bead-based enzyme-linked immunosorbent analysis

Yen-Heng Lin,^1, ^2, ^3, ^a) Ying-Ju Chen,¹ Chao-Sung Lai,^1, ^ii, ^b) Yi-Ting Chen,^iv Chien-Lun Chen,^v Jau-Vocal Yu,⁴ and Yu-Sun Chang^iv

Yen-Heng Lin

¹Department of Electronic Engineering, Chang Gung University, Taoyuan 333, Taiwan

ⁱⁱHealthy Aging Research Center, Chang Gung University, Taoyuan 333, Taiwan

ⁱⁱⁱGraduate Institute of Medical Mechatronics, Chang Gung University, Taoyuan 333, Taiwan

Ying-Ju Chen

¹Department of Electronic Engineering, Chang Gung University, Taoyuan 333, Taiwan

Chao-Sung Lai

^oneDepartment of Electronic Engineering, Chang Gung University, Taoyuan 333, Taiwan

²Healthy Aging Research Eye, Chang Gung Academy, Taoyuan 333, Taiwan

Yi-Ting Chen

^fourMolecular Medicine Inquiry Center, Chang Gung University, Taoyuan 333, Taiwan

Chien-Lun Chen

^fiveChang Gung Bioinformatics Center, Department of Urology, Chang Gung Memorial Hospital, Taoyuan 333, Taiwan

Jau-Song Yu

⁴Molecular Medicine Research Heart, Chang Gung Academy, Taoyuan 333, Taiwan

Yu-Sun Chang

⁴Molecular Medicine Enquiry Center, Chang Gung University, Taoyuan 333, Taiwan

Received 2013 Jan 9; Accepted 2013 February 26.

Abstract

This paper describes an integrated microfluidic chip that is capable of quickly and quantitatively measuring the concentration of a float cancer biomarker, apolipoprotein A1, in urine samples. All of the microfluidic components, including the fluid transport system, the micro-valve, and the micro-mixer, were driven by negative pressure, which simplifies the employ of the chip and facilitates commercialization. Magnetic beads were used every bit a solid back up for the primary antibody, which captured apolipoprotein A1 in patients' urine. Because of the iii-dimensional structure of the magnetic beads, the concentration range of the target that could be detected was every bit high as 2000 ng ml⁻¹. Considering this concentration is 100 times higher than that quantifiable using a 96-well plate with the same enzyme-linked immunosorbent assay (ELISA) kit, the dilution of the patient'southward urine can be avoided or greatly reduced. The limit of detection was determined to be approximately x ng ml⁻ⁱ, which is lower than the cutoff value for diagnosing float cancer (11.16 ng ml⁻ⁱ). When the values measured using the microfluidic chip were compared with those measured using conventional ELISA using a 96-well plate for five patients, the deviations were 0.9%, half-dozen.8%, 9.four%, 1.eight%, and 5.8%. The entire measurement time is 6-fold faster than that of conventional ELISA. This microfluidic device shows pregnant potential for point-of-care applications.

INTRODUCTION

Bladder cancer is a blazon of common urinary tract carcinoma that has a loftier recurrence rate and a poor prognosis.¹ ^, ^two If the abnormal tissue or tumor is identified early, handling and recovery may be easier. The standard method for the clinical detection of bladder cancer is cytology, which shows low sensitivity for depression-grade bladder cancers.^one ^, ² ^, ³ Cystoscopy is frequently used to examine and monitor patients for the recurrence or progression of this disease. Yet, this detection method is invasive and expensive.^four ^, ⁵ There accept been many attempts to develop an efficient, reliable, accurate, and noninvasive diagnostic procedure that can place bladder carcinoma patients. The quantitative measurement of urinary tumor biomarkers represents a practical method for the initial detection of tumors and for the monitoring of patients for recurrence because urine is in direct contact with tumor cells for this type of cancer and is attainable for clinical analysis. Recently, apolipoprotein A1 (APOA1) has been identified as a potential biomarker that can be used for the early diagnosis of bladder cancer.^ane ^, ⁶ ^, ⁷

Enzyme-linked immunosorbent assay (ELISA), a high-sensitivity technique, is the current standard method for the quantitative assay of a target protein in biological samples. The amounts of certain proteins in urine samples have been suggested to be reliable and quantitative indicators of bladder cancer.⁸ ELISAs have been used extensively in medical inquiry, clinical diagnostics, drug discovery, ecology monitoring, nutrient condom, and biodefense.⁹ Nevertheless, the conventional immunoassay conducted with a 96-well plate requires substantial labor, the consumption of expensive reagents, and precise technical functioning, making this type of assay inconvenient and impractical for point-of-intendance diagnosis.^ten ^, ¹¹ The principle of ELISA is to immobilize and detect an antigen–antibody complex. In conventional ELISA, the antibodies are immobilized on the surfaces of the wells in 96-well plates. During the assay process, the sample and reagent are dispensed by transmission pipetting or pipetting with a motorcar. Each incubation step is followed by repeated washing steps to remove unbound antibodies and nonspecific antigens. If the efficiency of the washing process is improved, the time required to complete the entire assay can be greatly shortened. In add-on, the analysis requires several dilution procedures to reduce the high concentration of the target antigen in urine samples. As a result, the assay requires a few hours or days for the liquid-treatment, washing, and incubation procedures. Furthermore, well-trained personnel are required to deport the entire protocol precisely.

Since its introduction, the lab-on-a-chip (LOC) organization, also known every bit the micro full analysis arrangement (μTAS), has been rapidly adopted to miniaturize several analytical assay systems that crave large amounts of sample or reagents and bulky experimental devices.¹² ^, ¹³ This technique offers many advantages over conventional techniques, such as brusque reaction times, less consumption of expensive chemical reagents, higher sensitivity, greater portability, and automatic performance.¹⁴ ^, ¹⁵ LOC systems are widely employed in analytical assays and clinical diagnosis. For example, electrophoresis was one of the earliest successes with microfluidics engineering science, and microfluidic electrophoresis devices have proven to be useful. A microfluidic electrophoresis bit with a polymerase chain reaction (PCR) Deoxyribonucleic acid amplification chamber was developed in 2000.¹⁶ In addition, the combination of microfluidics with menstruation cytometry provides a volume-efficient and college-precision method for determining the characteristics of suspended sample populations (due east.yard., cells, viruses, bacteria, yeast, droplets particles, and microbeads).¹⁷ This combination provides college speeds, smaller sizes, and lower costs than the conventional method. Moreover, the employ of a microfluidic-based electrospray fleck for mass spectrum assay has likewise been demonstrated. An enrichment column, a reversed-phase separation aqueduct, and a nanoelectrospray emitter were integrated into a single chip.¹⁸ This chip enhances the sample loading and selectivity of the mass spectrum system. Immunoassays tin besides be performed in microfluidic systems. For example, a miniaturized mosaic based on using a microfluidic network to pattern lines of antigens onto a flake surface has been presented.¹⁹ ^, ²⁰ The binding of target antibodies with their immobilized antigens on the surface results in a mosaic pattern with fluorescence labeling. This chip can exist used to perform dense, parallel, and self-consistent immunoassays with nanoliter quantities of reagents and incubation times of seconds to minutes. Furthermore, microfluidic engineering can provide a microenvironment for 3-dimensional prison cell cultures, which may ameliorate reverberate actual responses than 2D monolayer cultures can. This technique has been demonstrated for the empirical testing of tumor sensitivity to drugs, which may provide a more reliable predictive value for clinic diagnoses.²¹ Although microfluidic techniques provide various approaches for biological measurement, the sample pretreatment process for biosamples such as blood, saliva, and urine is still an important consequence affecting the general and practical use of these microfluidic devices.

Magnetic beads accept ofttimes been used in microfluidic immunoassays for several reasons. First, these beads offer greater surface area-to-volume ratios than the traditional 96-well plate, thus facilitating interactions between antigens and antibodies in small volumes. In addition, magnetic beads can be easily delivered using a flowing fluid and can be separated from the medium using a magnetic field. Furthermore, multiple antigen-antibody targets can exist included on a unmarried fleck if the magnetic beads are conjugated with different fluorophores. Magnetic chaplet allow sample pretreatment in microfluidic systems. For example, cell separation in microfluidic chips using a magnetophoresis approach has been proposed.²² Suspended cells labeled with magnetic chaplet in a microchannel are deflected by means of a magnetic field. Therefore, cells with different sizes or that are decorated with unlike numbers of magnetic beads are separated accordingly. The extraction of Deoxyribonucleic acid or RNA from a cell or virus is an important technique for molecular diagnosis. The utilise of magnetic beads facilitates the isolation of Dna or RNA from a cell or virus in a microfluidic device. Such a device had been developed to detect viral RNA with the help of magnetic beads.²³ ^, ²⁴ In addition, target viruses or leaner from a complex biosample can also be extracted directly using antibody-conjugated magnetic beads. Lien et al. purified and enriched dengue virus using magnetic beads. Next, nucleic acrid distension using a micro RT-PCR system was performed on the aforementioned chip.²⁵ In addition, magnetic beads with ion-exchange properties were used to simulate cells with peptides spring to the cell surface past electrostatic interactions. Past sequential treatment with dissimilar buffer weather condition using acoustophoresis-based microfluidic chips, the medium surrounding the beads can exist removed or replaced by some other buffer medium, which is important for the understanding of man biological science and the treatment of disease.²⁶ Using magnetic beads has benefits in biosample pretreatment when using LOC systems. Withal, the microfluidic devices proposed to date accept not been used for the diagnosis of float cancer in urine samples with a bead-based immunoassay using automatic fluid handling.

In this study, we developed a negative-pressure-driven microfluidic chip that can exist used to rapidly observe a bladder cancer biomarker with the help of magnetic bead-based ELISA. The use of merely vacuum forces as the driving forcefulness for fluid transport, the vibration of the micromixer, and the opening and closing of the microvalves simplifies the respective components of the chip. In addition, by integrating a syringe filter into the sample inlet, the centrifugation step in the pretreatment of urine samples is no longer necessary. Five clinical urine samples were used to evaluate the use of the chip. The proposed chip is a feasible sample-in and reply-out device that can quantitate a potential bladder cancer biomarker.

MATERIALS AND METHODS

Fleck design and experimental setup

Fig. 1a is an exploded view of the chip. The chip consists of four poly(dimethylsiloxane) (PDMS) layers: the summit cover layer, the air chamber layer of the normally closed valve, the fluid aqueduct layer, and the micromixer air chamber layer at the bottom. The channel depths of the normally closed valve air chamber layer and the fluidic aqueduct layer were both 500μm. The bottom micromixer air sleeping accommodation layer was constructed using a thick PDMS layer with a channel depth of 1.5 mm. There were five reservoirs that contained primary antibody-coated beads, secondary antibodies, antigens, enzymes, or the washing buffer, as shown in Fig. 1b. Each reservoir was connected to the central sleeping room through a usually airtight valve. The cardinal chamber was equipped with a micromixer and served as a reaction bedchamber with a volume of xiv.5μl. Notation that all the driving forces—including those powering fluid ship, the micromixer, and the normally closed valves—were based on negative force per unit area to simplify the experimental setup. Other types of negative-pressure-driven microfluidic components can be found in the literature.²⁷ ^, ²⁸ ^, ²⁹ The normally closed valve and micromixer were controlled and driven using a vacuum pump connected to an electromagnetic valve (EMV), and fluid transport was accomplished using a syringe pump (Legato 180, KD Scientific Inc., Holliston, MA) connected to the waste outlet channel that was driven with a withdrawing force by opening the inlet to atmospheric pressure. Using a personal computer, control of the fluid in the microchannel can be performed automatically. For example, when the washing process was conducted, the valves betwixt the washing buffer reservoir and the reaction chamber and between the reaction chamber and the waste aqueduct were opened, and the syringe pump withdrew the remaining washing buffer to consummate the washing process. This system enabled five reagents to be delivered sequentially into the reaction chamber via negative-pressure-driven fluid transport and commonly airtight valves. The micromixer was composed of ii air chambers with thin PDMS membranes, and the ii air chambers were continued with a serpentine-shaped air channel. The serpentine-shaped connecting air channel was designed to filibuster the airflow from one micromixer air chamber to the other. Therefore, the two thin PDMS membranes higher up the air chamber were deflected sequentially, leading to the swirling of the fluid in the reaction sleeping accommodation later negative air pressure was practical to the air inlet (run into details in Fig. 4a). Thus, the reagents and sample tin can exist incubated with enhanced mixing efficiency. Fig. 1c shows a photograph of the integrated chip in which the normally closed valves are presented in green, the fluidic sleeping room is in blueish, and the micromixer is in carmine. A commercially available syringe filter with a pore size of 0.22μm was continued to the antigen loading chamber to filter out unwanted debris in the urine sample. The dimensions of the chip were twoscore mm in length by 40 mm in width.

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(a) Exploded view of the proposed microfluidic bit. 4 layers of PDMS were used to construct the negative-pressure-driven microfluidic fleck. The bottom layer was an air bedchamber layer for the actuation of a pneumatic micromixer, the third layer was a fluidic aqueduct layer, the 2nd layer was another air chamber layer for the actuation of the commonly airtight valves, and the summit layer was a flat PDMS layer that sealed the air bedroom. (b) The chip was equipped with a micromixer incorporated into the reaction chamber at the center of the flake and with five reservoirs (4 for sample loading and one for a wash buffer). At that place were six usually airtight valves located betwixt the reservoirs and the reaction bedchamber. A syringe filter was integrated into the antigen-loading chamber to filter out any debris in the urine samples. The driving strength for fluid motility relied on a suction force provided from the waste outlet. All of the liquid treatment tin be performed with the aid of the integrated microfluidic components. (c) Photo of the assembled microfluidic chip with dimensions of xl mm × forty mm.

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Blueprint and characterization of the suction force-driven micromixer. (a) Cantankerous-sectional view of the micromixer. The mixer has 2 air chambers connected by a narrow air channel to filibuster the movement of air from i bedchamber to the other. When pulses of suction force were practical, the fluid higher up the PDMS membrane vibrated, thus enhancing the mixing efficiency. (b) Series of photographs of the mixer actuated for a period of 30 south. The ink and DI water were mixed together gradually. (c) Label of the mixing alphabetize of the micromixer with unlike negative air pressures and driving frequencies.

Experimental procedure

In this written report, the detection of a protein, APOA1, was used to demonstrate the feasibility of using the proposed bit for cancer detection. This protein is a highly relevant protein associated with bladder cancer in urine samples. The concentrations of APOA1 were adamant using a bead-based approach with a sandwich ELISA kit (Mabtech, Nacka Strand, Sweden). Fig. two shows a schematic illustrating the principle upon which the bead-based ELISA in the microfluidic chip is based. First, the chief antibiotic-coated magnetic beads and the antigen were incubated together in the reaction chamber to allow the beads to specifically capture the antigen, APOA1. The protocol for conjugating the antibiotic on the magnetic beads is described in the supplemental data section.^thirty After the antigen and beads were co-incubated, the beads were held in the chamber using an external magnet, and the nonspecific antigens were washed out of the reaction chamber (Figs. 2a, 2b). Serially diluted, purified APOA1 was used as the calibration standard. The APOA1 signals for the clinical urine samples were measured, and the standard bend was used to calculate the APOA1 concentration. Biotinylated secondary antibodies were allowed to collaborate with the antigens, and then unbound antibodies were washed away (Figs. 2c, 2d). Next, streptavidin-enzyme complexes were allowed to collaborate with the secondary antibody (Fig. 2e). Finally, the substrate was added to the reaction chamber and allowed to react with the enzyme. And then, the reagents were suspended to measure the optical density at 405 nm (Fig. 2f).

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Illustration of the working principles behind the bead-based ELISA using the microfluidic chip. (a) A urine sample containing the target protein and antibiotic-coated magnetic beads was introduced into the chip. (b) After the incubation procedure, the target protein was captured, and the unwanted protein was washed out. (c) and (d) The secondary antibody was bound to the antigen, and the backlog antibody was washed out. (e) and (f) The enzyme was linked to the secondary antibiotic, and after the excess enzyme was washed out, a substrate was used to quantitatively mensurate the target protein.

The scrap blueprint enabled the reagents to be sequentially and automatically introduced into the reaction chamber. The detailed operation procedure for the measurement of APOA1 concentrations using the proposed chip is as follows. As shown in Fig. 1b, initially, the valves that connect the bead-loading reservoir to the reaction chamber and the reaction chamber to the waste aqueduct were opened, and 20μl of antibody-coated magnetic beads (ten^ix beads ml⁻¹) was loaded into the reaction sleeping accommodation using the suction force from the outlet channel. Annotation that before the magnetic beads were introduced, an external magnet (approximately 300 Gauss) was placed under the reaction chamber to gear up the beads in the reaction chamber. The solution in which the beads had been suspended was then suctioned out through the outlet aqueduct. Then, xxμl of standard APOA1 solution or a urine sample was loaded into the reaction chamber using the same command procedure. Note that the book of the used reagent was determined by the size of the reaction sleeping accommodation, i.e., 14.fiveμfifty. After the normally closed valves were closed, the excess reagent either left via the microchannel or was suctioned out. Next, the sample and magnetic beads were mixed for five min using a 2-membrane-type micromixer. Later the incubation period, the chaplet were again held in identify using the magnet, and the fluid was withdrawn from the waste channel using a syringe pump. During the washing process, the valves that connect the washing buffer reservoir to the reaction chamber and the reaction sleeping room to the waste material channel were opened. Thus, the washing buffer flowed into the reaction chamber for 30 s. Next, the valves were closed, and the washing buffer was agitated for one min using the micromixer to remove the unbound antibody. This procedure was repeated three times for a single washing step. Other reagents, including the secondary antibody and the enzyme, were transferred into the reaction sleeping accommodation and mixed sequentially. Similarly, each incubation pace lasted 5 min, and the washing stride was repeated iii times for 1.five min. Finally, the substrate was loaded into the reaction sleeping room from the bead-loading reservoir and was reacted with the enzyme for ten min. And then, the reagents were suspended for optical measurement.

Scrap fabrication

The chip was made using standard PDMS photolithography, micromachining, and replication techniques. Poly(methyl methacrylate) (PMMA) primary molds for the microstructures were first formed using a CNC machine (EGX-400, Roland Inc., Nippon) equipped with a 0.5 mm drill bit. A silicone elastomer and an elastomer curing agent (Sylgard 184A and 184B, Sil-More Industrial Ltd., USA) were mixed in a x:1 ratio and poured onto the PMMA mold. To form a 100-μthou-thick PDMS membrane to drive the microvalve and the micromixer, the microvalve air chamber layer and the fluidic aqueduct layer were fabricated by spin blanket with three.half-dozen g and iv g of PDMS, respectively, at 300 rpm for xv s. A 600-μm-thick PDMS layer was formed with a 100-μthou-thick PDMS membrane located on the microstructure of the master mold. After curing at 70 °C for 1 h, the solidified PDMS was peeled off of the master mold, and the holes for injecting air or fluid were drilled with a syringe needle. Subsequently, the iv-layer PDMS construction was bonded using oxygen plasma treatment (HARRICK PLASMA, Ithaca, NY, USA).

Immunoassay reagents and urine sample training

Phosphate-buffered saline (PBS), bovine serum albumin (BSA), and Tween-twenty surfactant were purchased from Sigma-Aldrich Co. Superparamagnetic beads (M-270 Epoxy, Dynabeads, Invitrogen) were coated with a surface epoxy grouping. These beads were used as a solid support to immobilize specific antibodies. In this study, all urine samples were collected at Chang Gung Memorial Infirmary, Taoyuan, Taiwan (a hospital affiliated with Chang Gung University). First, morning urine samples were collected and treated with one mM sodium azide (Fluka, Switzerland) as a bacteriostatic reagent. A hernia patient (due north = 1) was used every bit a control. The nerveless urine samples were stored at −20 °C for subsequent processing.

RESULTS AND DISCUSSION

Characterization of the microfluidic components

Fig. 3a illustrates the working principles of the normally closed valve. As shown in the lateral view, the normally closed valve was composed of an air sleeping room, a thin PDMS membrane layer, and a pillar structure in the microchannel. The PDMS membrane and the pillar structure were not bonded together. When negative air force per unit area from a vacuum was applied to the air sleeping accommodation, the PDMS membrane was plain-featured and actuated, assuasive the fluid to flow through the microvalve. After the vacuum was removed, the PDMS membrane recovered, and thus, the fluid menses was blocked by the valve. Fig. 3b shows a valve in the open and the closed position. The commonly closed valve was characterized using a constant gravity-driven flow of deionized (DI) water (xiv ml with a height of 10.5 cm). The time required for the start v ml of DI water to flow through the scrap at different practical vacuum pressures was recorded. The catamenia charge per unit was calculated using these data. By adjusting the vacuum force per unit area, the period rate of the fluid that passed through the valve was varied. No fluid passed through the valve when the applied pressure was less than −4.2 kPa, i.e., the threshold vacuum pressure was −4.2 kPa for valve actuation. The flow charge per unit increased from 16.4μl s⁻¹ to 312.fiveμfifty due south⁻¹ when the actuation force per unit area was inverse from −5 kPa to −45 kPa (as shown in Fig. 3c). When the practical pressure was greater than −45 kPa, the normally airtight valve was virtually fully open.

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Pattern and label of the normally closed valve. (a) Working principle of the valve. When the air in the sleeping room was suctioned out, the PDMS membrane was deflected, and the fluid could pass through the valve. (b) Photograph of the valve in the open and closed positions. (c) The DI water started flowing through the normally closed valve when the actuation pressure surpassed −iv.2 kPa. The normally closed valve was nearly fully open up when the applied pressure was greater than −45 kPa.

The ii-membrane-type micromixer was designed to mix the urine sample and reagents during incubation. Fig. 4a shows the cross-sectional view of the micromixer. The two membranes were connected by a serpentine-shaped connecting channel, and the air inlet was linked to a vacuum forcefulness. Past setting the driving frequency of an EMV and using a programmable computer, the ii thin membranes were vibrated at a specific frequency and in a specific sequence to raise the mixing efficiency during incubation. To optimize the mixing efficiency of the two-membrane-type micromixer, various air force per unit area and practical frequency conditions were investigated. To make up one's mind the mixing efficiency of the micromixer, 2μl of blue ink and 12.vμfifty of DI h2o were loaded into the reaction chamber to evaluate the concentration distribution along a cross section of the reaction chamber. A mixing index was defined to quantify the mixing profile:³¹

$ρ (A) = (1 - \frac{\int_{A} | C^{+} - C_{\infty}^{+} | d A}{\int_{A} | C_{0}^{+} - C_{\infty}^{+} | d A}) \times 100 %,$

(i)

where ρ(A) is the mixing index of the normalized concentration (C⁺ ) distributed within the sample mixing unit (A), $C_{0}^{+}$ is the initial status in the unmixed state, and $C_{\infty}^{+}$ is the completely mixed state of the normalized concentration (=0.5). Fig. 4b shows the concentration distribution of the fluid in the mixing chamber for mixing periods of 0, 2, iv, 6, 10, and thirty south. After 30 southward, the mixing alphabetize increased from 15 to 95% at a driving frequency of 35 Hz and an air pressure of −80 kPa. The initial mixing index of 15% was the event of the molecular diffusivity in the fluid. The time required to achieve a 100% mixing alphabetize was calculated to be approximately 60 s. However, this length of time was not sufficient for sample incubation. In practise, each incubation footstep lasted v min.

Fig. 4c shows the mixing index of the micromixer under applied air pressures of −60, −70, and −80 kPa and driving frequencies of 10, 15, 20, 25, xxx, and 35 Hz after thirty s. A higher mixing index was achieved with college applied negative air pressures and college practical frequencies. For an air pressure level of −80 kPa, the mixing index was 95% at an applied frequency of 35 Hz for 30 due south. At the practical air pressures of −threescore and −70 kPa, the 100-μm PDMS membrane could not be deformed completely. Therefore, the mixing index could not accomplish values as high as that for an applied pressure level of −80 kPa. The suction-force-driven pneumatic micromixer greatly enhanced the mixing efficiency and shortened the incubation fourth dimension of the dewdrop-based ELISA for measurement of the APOA1 concentration. However, fatigue is a problem for PDMS membranes during on-off cycling, although the literature indicates that a PDMS membrane can perform more than 4 × x⁶ actuations without fatigue failure.³² Therefore, we used each chip three to five times to finish one sample measurement, and and so used some other chip for the next sample measurement to make sure that the procedure was performed past a mechanically stable PDMS membrane.

The maximum mixing efficiency was observed at a driving frequency of 35 Hz with a suction force of −80 kPa.

Detection of the float cancer biomarker APOA1

Prior to detecting the patients' samples, the incubation time for the flake was determined. Supplemental Fig. S1³⁰ shows the relationship betwixt the detection signal and the mixing time for three concentrations of APOA1 within 20 min. Note that the mixing time represents every incubation step in each assay. It can exist observed that the detection signal was increased with the increased mixing fourth dimension. Notably, the betoken increased speedily before the incubation time reached 5 min. From time menstruum of 5 min to xx min, the point increased 22%, 28%, and 26% with APOPA1 concentrations of thousand, 500, and 200 ng ml⁻¹, respectively. We, therefore, chose 5 min for each incubation period, which was a trade-off between time and signal intensity. The detection limit of v min (10 ng ml⁻¹) is sufficiently smaller than the cut-off valve (11.16 ng ml⁻¹) for bladder cancer diagnosis using APOA1 as a biomarker. Therefore, each incubation pace lasted 5 min. The detection range of APOA1 was measured using the proposed system. APOA1 standards with concentrations ranging from 0 to 10000 ng ml⁻¹ were prepared using proper dilution ratios. As shown in Fig. 5a, when the concentration of APOA1 was less than 2000 ng ml⁻¹, the optical point increased proportionally to the APOA1 concentration with a linearity of 0.996. The calibration curve of APOA1 was calculated for repeated experiments (n = v). When the concentration of APOA1 exceeded 2000 ng ml⁻ⁱ, the optical point exhibited a nonlinear plateau. This finding revealed that the detection range for the APOA1 concentration tin can reach 2000 ng ml^−one when using 20μfifty of magnetic beads (1 × 10^nine beads ml⁻¹). In comparing, the detection range using the tradition method with a 96-well ELISA plate with the same ELISA kit was 0.ii to xx ng ml⁻¹. Thus, the traditional plate-based ELISA requires several repeated dilutions to mensurate the high concentration of APOA1 in clinical urine samples, and this requirement results in burdensome and complicated assay procedures. The maximum concentration of APOA1 in urine is approximately 9000 ng ml^−ane.⁶ Therefore, at well-nigh, one 10-fold dilution is necessary to analyze unknown samples using the proposed chip. With the profoundly enhanced detection range of ELISA using the proposed method, the procedure of diluting samples can be omitted or reduced, which is beneficial when needing to clarify a large number of patient samples. Because the magnetic beads have a larger surface surface area-to-volume ratio than the 96-well plate, the beads increase the efficiency of interactions between antigens and antibodies, further enhancing the detection range of the immunoassays and shortening the measurement time. Although the cutoff value for diagnosing bladder cancer was xi.16 ng ml⁻¹, we also tried to decide the relationship between the stage of the bladder cancer and the concentration of APOA1.^half-dozen If such a relationship can be identified, the proposed chip can exist used non only for the early detection but as well for phase differentiation of bladder cancer. To make up one's mind the detection limit, an APOA1 concentration range from 0 to 20 ng ml⁻¹ was used. Figure 5b shows the background noise and the detection signal for the low concentrations of APOA1. The detection limit of APOA1 on the chip was determined to be approximately x ng ml^−ane. The cutoff value for diagnosing bladder cancer using the APOA1 poly peptide has been reported to exist 11.16 ng ml⁻¹ (north = 126, 94.six% sensitivity and 92.0% specificity).⁶ Thus, the detection limit of 10 ng mL⁻ⁱ is sufficient to diagnose bladder cancer with a urine sample. The fleck requires but 5 min for each incubation procedure, which is less than the 1 h required for the 96-well plate on a shaker. Thus, the full measurement fourth dimension with the bit is profoundly reduced, totaling approximately 40 min, which is much shorter than the time needed to run a conventional plate-based ELISA (more than than 4 h). In some other bead-based microfluidic analysis approach using surface tension valves for fluidic control,³³ the measurement of the concentration of the biomarker in that chip was performed off-flake, only in the proposed chip, it was performed in the same chip, which eliminates the need for manual operation and facilitates automated control of the measurement process.

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The detection power of the proposed chip. (a) The detection range was upwardly to 2000 ng ml⁻¹ with 20μfifty of magnetic beads (1 × 10⁹ chaplet ml⁻¹), and the detection linearity was 0.996 in this range. (b) The detection limit was approximately 10 ng ml⁻¹ for APOA1 for the scrap.

Detection of APOA1 in clinical urine samples

To evaluate the feasibility of using the proposed chip for the assay of clinical urine samples, patient urine samples collected at the Chang Gung Memorial Hospital were used. The experimental results were then compared with the results of conventional ELISA conducted in a 96-well plate. One normal control urine sample from a hernia patient and four samples from bladder cancer patients were tested. The concentrations of APOA1 in the urine samples (1, 2, 3, 4, and v) were 4.24, 207.3, 826.3, 1038.7, and 3754.7 ng ml⁻¹, respectively. With the exception of sample 5, the samples were measured in their native undiluted grade. The concentration of APOA1 in sample number 5 was diluted x-fold prior to analysis. Each sample was assessed three times (n = three), and the mean value was calculated. Because clinical urine samples may contain debris, a syringe filter was integrated into the inlet of the flake to eliminate unwanted debris from the urine sample. Fig. half dozen shows the experimental results for the quantitative detection of APOA1 using the proposed fleck and the traditional ELISA method. The variations betwixt the 2 sets of results were found to exist 0.9%, 6.8%, ix.4%, 1.8%, and 5.8%. Note that the values represented by the error bars are the maximum/minimum values of the experimental data. The variation between the two sets of results may exist the result of several factors, such as the absorption of protein to the PDMS surface or stiction between the magnetic chaplet and the PDMS surface, which tin cause dewdrop loss during each incubation process. The experimental data (Fig. 6) prove that most measurement results for the APOA1 concentration that were obtained with the proposed bit were a fiddling lower than the results from the ELISA plate. In addition, the absorption of protein on PDMS surfaces is a well-known upshot when PDMS is used in microfluidic chips.³⁴ ^, ³⁵ Therefore, we infer that the absorption of poly peptide on the PDMS surface is ane of the factors that acquired the divergence in detection. Furthermore, observations fabricated during the experimental procedure indicated that there were some magnetic beads attached to the microchannel surface that could non participate in the reaction. This might be the other reason why the detected concentration in the flake was lower than in the ELISA plate. However, in general, the results confirm that the detection functioning of ELISA using the proposed device is broadly comparable to that of traditional ELISA only that approximately half-dozen-fold less time is required for the chip-based analysis.

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Comparison of the results for v patient samples obtained using the proposed chip and using traditional 96-well ELISA. The differences in the results were 0.9%, half-dozen.viii%, nine.4%, 1.8%, and 5.eight% for samples 1, 2, 3, 4, and five, respectively.

Conclusion

We evaluated the use of a microfluidic chip equipped with a negative-pressure-driven microvalve, micromixer, and fluid transport system for bead-based ELISA, and we used this form of ELISA to successfully quantitatively measure the bladder cancer biomarker APOA1 in human urine. The microvalve was determined to take an applied air force per unit area threshold of −4.2 kPa to open the valve, and the optimal mixing efficiency of the suction-blazon micromixer was found to occur at an air pressure of −80 kPa and a frequency of 35 Hz. Because magnetic beads were used equally the solid support for the primary antibiotic, the surface expanse-to-volume ratio for the interaction between the reagent and sample was greatly increased compared with the ratio for the plate-based ELISA. Therefore, the concentration range for measurement was greatly enhanced, and the measurement time was besides reduced. Five human urine samples were analyzed using the proposed bit, and the results were compared with those from the 96-well plate-based ELISA exam. The maximum variation between the two sets of results was plant to be ix.4%, for patient number three. In general, the operation of the proposed scrap was comparable to that of 96-well plate-based ELISA. The proposed microfluidic device is a promising tool for the bespeak-of-care diagnosis of float cancer.

ACKNOWLEDGMENTS

This work was supported past grants to Chang Gung University from the Ministry building of Education (EMRPD1A0761) of Taiwan, Republic of China; Chang Gung Academy (UERPD2A0051, UERPD2B0091); and the National Scientific discipline Council of Taiwan, Republic of Mainland china (100-2221-E-182-021-MY3).

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Source: https://www.ncbi.nlm.nih.gov/pmc/articles/PMC3606202/