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Microfluid Nanofluid (2017) 21:74 DOI 10.1007/s10404-017-1910-y

RESEARCH PAPER

A magnetophoresis‑based microfluidic detection platform under a static‑fluid environment

Younggeun Jo1 · Young Ki Hahn2 · Je‑Kyun Park1

Received: 7 November 2016 / Accepted: 25 March 2017 © Springer-Verlag Berlin Heidelberg 2017

semiquantitative detection of labeled particles without any complex electrical devices and could thus be used as a port-able detection platform.

Keywords Easy magnetophoretic assay · Pumpless platform · Magnetic nanoparticles · Portable device · Static-fluid environment

1 Introduction

Microfluidics has widely been adapted as a powerful approach for the detection of biomolecules such as proteins and nucleic acids from a biological sample as it allows for efficient and rapid detection of target molecules from a small sample. In particular, microbead-based immunoassay has been actively developed using microfluidic technolo-gies because microbeads offer several advantages over a planar substrate, including a large surface-to-volume ratio, availability of the multiplex assay via various encoding methods such as fluorescence (Gao et al. 2013; Zhang et al. 2016), color (Hahn et al. 2009) and the surface pattern of microparticles (Lee et al. 2010), and the ability to manipu-late in microchannels by external forces. The manipulation of microbeads in microchannels provides some benefits from sample preparation steps, including trapping, incuba-tion, washing, detaching, and transporting of the microbe-ads in microchannels. To date, various technologies have been developed using optoelectronic tweezer (Han et al. 2013), magnetophoresis (Lin and Peng 2015; Sasso et al. 2012), dielectrophoresis (Emaminejad et al. 2014), and acoustophoresis (Tenje et al. 2015) for the manipulation of microbeads in microchannels. Meanwhile, the forces acting on microbeads have also been utilized for quantifying the labels on the microbeads (Han et al. 2013).

Abstract Microfluidic cell separations and immunoassays exploit a dynamic flow environment by electrical pumps to manipulate fluids containing biomolecules and microbe-ads. In particular, the magnetophoresis-based microfluid-ics requires a delicate flow control of pumps because the flow rate affects the result sensitively. Consequently, the dynamic flow environment requiring pumps prevents the magnetophoresis-based microfluidics from populariza-tion and miniaturization. Herein, we present a magneto-phoresis-based microfluidic platform under a static-fluid environment for the detection of microbeads labeled with magnetic nanoparticles (MNPs) by simple manual opera-tion of fluids. To overcome the residual flow caused by the manual operation, we designed a microfluidic device hav-ing a pair of microchannels; one for detecting the target and the other for a reference. The deviations due to the residual flow were corrected by comparing the difference between the mean velocities of microbeads in each microchannel where microbeads labeled with five different concentra-tions of MNPs could be classified. On the basis of the con-venience and portability of magnetophoresis under a static-fluidic environment, this new microfluidic platform enabled

This article is part of the topical collection “2016 International Conference of Microfluidics, Nanofluidics and Lab-on-a-Chip, Dalian, China” guest edited by Chun Yang, Carolyn Ren and Xiangchun Xuan.

* Je-Kyun Park jekyun@kaist.ac.kr

1 Department of Bio and Brain Engineering, Korea Advanced Institute of Science and Technology (KAIST), 291 Daehak-ro, Yuseong-gu, Daejeon 34141, Republic of Korea

2 Samsung Electronics, 4 Seocho-daero 74-gil, Seocho-gu, Seoul 06620, Republic of Korea

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Magnetophoresis is a technology exploiting the move-ment of particles (or cells) under inhomogeneous magnetic field by the difference in magnetic susceptibility between the particle and the surrounding solution. The magnetopho-retic device, more specifically based on a microchannel, has several advantages such as easy fabrication, low energy consumption, and convenient integration with a permanent magnet, while other technologies such as dielectrophoresis, acoustophoresis, and optoelectronic tweezers require com-plex fabrication and various electrical equipments. Thus, the magnetophoretic manipulation technology has been easily utilized in microfluidics for trapping (Ino et al. 2008; Lund-Olesen et al. 2007; Saliba et al. 2010) or separation (Adams et al. 2008; Jo et al. 2016; Robert et al. 2011; Zhu et al. 2010) of cells, microbeads, and droplets. In general, magnetic beads have been used as a substrate or carrier for labeled compounds in magnetophoretic applications, and thus, their detection methods exploit fluorescence (Rissin et al. 2010; Tennico et al. 2010), electrochemistry (God-ino et al. 2009; Tang et al. 2011), and surface-enhanced Raman spectroscopy (Chon et al. 2010) on the basis of labeled compounds. However, these detection technologies also require complex electrical equipment for reading the signals as well as laborious fabrication and experimental setup. To overcome this limitation, magnetophoresis-based microfluidic technologies using magnetic nanoparticles (MNPs) as a label were previously proposed by our groups (Hahn et al. 2007, 2009; Hahn and Park 2011; Jin et al. 2009). Unlike other works, we exploited the movement of the MNP-labeled microbeads by only a small permanent magnet as a detection signal of proteins. Although these magnetophoretic approaches have an obvious strength in detection aspects, a limitation still remains to be improved. Similar to other microfluidic technologies, magnetophore-sis-based microfluidics has exploited a dynamic flow envi-ronment in a microchannel, which needs additional control of electrical syringe pumps to manipulate fluids containing magnetic particles and microbeads or cells labeled with MNPs. In magnetophoresis-based microfluidics, a dynamic flow in a microchannel has been typically utilized for align-ment of microbeads with the same distance from a magnet because the magnetic field strength decreases according to the distance. For maintaining the dynamic flow environ-ment in a microchannel, a delicate and skilled technique is needed for the users to construct a desired flow envi-ronment such as the velocity of the microbeads, the ratio of the inlet flows, and the alignment of the microbeads, which could affect the result. Meanwhile, although iso-magnetophoresis-based microfluidics showed a good per-formance for protein detection (Hahn and Park 2011), this approach also requires a gradient of magnetic susceptibil-ity—the gradient profile sensitively affects the results—in a microchannel which is difficult to maintain throughout the

experiment. Despite several advantages of magnetophore-sis, the difficulty in controlling the dynamic flow environ-ment in microchannels prevents the magnetophoresis-based detection technology from popularization. For convenience and availability of magnetophoresis, the magnetophoretic device has to be simpler and more user-friendly.

Herein, we present a magnetophoresis-based micro-fluidic platform under a static-fluid environment to detect MNP-labeled microbeads. For magnetophoresis under a static-fluid environment, unique microchannels which could align injected microbeads without pumps are designed for simple and easy operation. The amount of MNPs bound to microbeads is detected by measuring the magnetophoretic velocity of the microbeads under a static-fluid environment, without delicate control of pumps.

2 Concept and principle

2.1 Concept

Considering polystyrene microbeads having diamagnetic property in paramagnetic solution under an inhomoge-neous magnetic field, the microbeads are moved away from the magnet field by a magnetic repulsive force. On the other hand, MNP-labeled microbeads, in which their magnetic susceptibilities are larger than those of the bare microbeads, are slowly moved away from the magnet field than the bare microbeads due to the additional mag-netic attractive force acting on the MNPs (Fig. 1). The movements of two types of microbeads, in a static envi-ronment, are only affected by the magnetic field. How-ever, the movement of microbeads under unpredictable flow in a microchannel is unsettled because the move-ment is no longer under the influence of only magnetic force. Consequently, the movement of microbeads shows large errors in the results due to unpredictable flow. These signal noises by unpredictable flow could be eliminated by comparing the velocity between the bare microbeads and the MNP-labeled microbeads located in the same environment.

Whole experimental procedures are designed to be free from the dependency of a dynamic flow environ-ment which requires delicate and relatively difficult con-trol of electrical pumps to the end user. For convenience of the experimental procedure, microbeads and solution are injected into each inlet of microchannels via drop-ping (Fig. 2). The injected microbeads are pulled by sim-ple handling of connected syringes and aligned in front of uniquely designed micropillar structures in order for the microbeads to have the same environment of the mag-netic field. After alignment of the microbeads, a port-able magnet is placed at the assigned position near the

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poly(dimethylsiloxane) (PDMS) chip. After then, the movement of the microbeads is observed to detect labeled MNPs. However, a rapid change in the pressure of a microchannel caused by manual operation of syringes gives rise to an uncontrollable flow (we called it residual flow) even after stopping the operation of the syringe. This uncontrollable residual flow, which causes errors in the results, is supposed to be caused by several reasons such as the hydraulic resistance of microchannels and the internal compressive materials such as air bubbles in microchannels (e.g., Tygon tubes and syringes). For the

elimination of air bubbles in a microchannel, a high pres-sure has been typically utilized by electrical pumps in a microfluidic device. Air bubbles in a microchannel are squeezed out through an air-permeable PDMS chip by the high pressure. However, this method could not elim-inate air bubbles in Tygon tubes and syringes. To over-come the residual flow, we designed two microchannels as a single unit to have the same environment. For the elimination of errors by residual flow, we also exploited a comparison of the velocity of the microbeads between two microchannels.

Fig. 1 Schematics of a magne-tophoresis-based microfluidic platform under a static-fluid environment

Fig. 2 a Design of a micro-fluidic device for the mag-netophoretic assay under a static-fluid environment. *NC negative control. b The device consists of a permanent magnet, a nickel structure, a manually operated syringe, and a PDMS microfluidic chip. c Micropillar array near the nickel structure is used for alignment of injected microbeads

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2.2 Principle

The magnetic repulsive force (Fm_r) and the attractive force (Fm_a) acting on an MNP-labeled microbead in par-amagnetic solution are given by Eqs. (1, 2):

where Vb is the volume of a bare microbead, ∆χb is the difference in magnetic susceptibility between a bare microbead and a paramagnetic solution (m3 kg−1), ∇B is the gradient of the magnetic flux density (T/m), μ0 is the vacuum permeability (N/A2), Nm is the number of labeled MNPs, Vm is the volume of an MNP, and ∆χm is the dif-ference of the magnetic susceptibility between an MNP and a paramagnetic solution. Considering the movement of the microbeads by magnetic force and the unpredicta-ble residual flow (vr) in the opposite direction, the Stokes’ drag force (Fd) is represented by the following Eq. (3):

where vsample is the velocity of an MNP-labeled micro-bead, vr is the velocity of residual flow, Rb is the radius of the microbead, and η is the viscosity of the paramag-netic solution. Considering Eqs. (1)–(4), the velocity of an MNP-labeled microbead (vsample) is represented by the following Eq. (5):

This velocity of the MNP-labeled microbead is affected by an uncontrollable residual flow, resulting in poor repeatability. This limitation should be overcome by comparison with the movement of a bare microbead in the same environment. Considering a bare microbead in the same environment, the magnetic repulsive force is the only force acting on a bare microbead unlike an MNP-labeled microbead:

Considering Eqs. (1), (6), and (7), the velocity of a bare microbead (vrefer) is affected by the magnetic repul-sive force and the residual flow:

(1)Fm_r =Vb�χb∇B

2

µ0

(2)Fm_a = −NmVm�χm∇B

2

2µ0

(3)Fd = −6πRbη�v = −6πRbη(

vsample + vr

)

(4)Fm_r − Fm_a = −Fd

(5)vsample =2Vb�χb∇B

2 − NmVm�χm∇B2

12πRbηµ0

− vr.

(6)Fd = −6πRbη�v = −6πRbη(vrefer + vr)

(7)Fm_r = −Fd .

(8)vrefer =2Vb�χb∇B

2

12πRbηµ0

− vr.

The velocities of both an MNP-labeled microbead and a bare microbead are affected by the same residual flow. Accordingly, the effect of residual flow could be removed by the velocity difference between an MNP-labeled micro-bead and a bare microbead:

This velocity difference is only affected by the magnetic attractive force caused by labeled MNPs. This comparison of the velocity between reference and sample microbeads overcomes the limitation caused by manual operation for pumpless magnetophoretic assays.

3 Materials and method

3.1 Materials

A permanent magnet (NdFeB35; w × l × h = 25 mm × 50 mm × 10 mm) was purchased from Magtopia (Gumi, Korea). PDMS base monomer and curing agent (Sylgard 184) were obtained from Dow Corning (Midland, MI, USA). Gadolinium–diethylenetriamine pentaacetic acid (Gd-DTPA), bovine serum albumin (BSA), Teflon AF, and FC-40 were purchased from Sigma-Aldrich (St. Louis, MO, USA). Polystyrene microbeads with a diameter of 20 μm were obtained from Polysciences, Inc. (Warrington, PA, USA). Rhodamine B-marked polystyrene microbe-ads with a diameter of 2 μm were purchased from Fluka Chemie (Buchs, Switzerland). Streptavidin-functionalized MNPs with a diameter of 130 nm were obtained from Micromod (Rostock, Germany).

3.2 Device fabrication

A microfluidic device consists of a PDMS chip and a nickel microstructure. The PDMS chip was fabricated by soft lithography technique. First, a negative photore-sist, SU-8 2050 (Newton, MA, USA), was coated on a Si wafer to reach a thickness of 100 μm using a spin coater at 1700 rpm. The photoresist-coated wafer was baked at 65 °C for 4 min and at 95 °C for 15 min. After the baking step, the wafer was exposed to ultraviolet (UV) light (230 mJ/cm2) via a photomask for the patterning of photoresist. For post-baking, the UV-exposed photoresist was placed in an oven for 4 min at 65 °C and 10 min at 95 °C and a baked photoresist was developed for 12 min. The mas-ter mold was silanized to promote the release of a PDMS replica. A PDMS base monomer and curing agent were mixed at a 10:1 ratio and degassed. Then, it was casted on the master mold and cured at 90 °C for 50 min. After bak-ing, the PDMS chip was peeled off the master. The PDMS

(9)�v = vrefer − vsample =NmVm�χm∇B

2

12πRbηµ0

.

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chip was bonded to glass slide. A nickel microstructure was fabricated by a nickel electroplating method. First, chro-mium (Cr) and gold (Au) were deposited on a glass wafer by metal sputtering for adhesion of nickel. Then, a nega-tive photoresist, THB-151N, was coated on Cr/Au seed layer and pattered by a standard lithography technique. After nickel electroplating, the nickel was polished to a thicknesses of 100 μm using a chemical mechanical pol-ishing method. Then, the photoresist was removed using a THB-S1 stripper (JSR Corp., Tokyo, Japan) for about 2 h at 60 °C. The nickel microstructure (w × l × h = 2 mm × 10 mm × 100 μm) was detached and positioned near the microchannel. This nickel structure was easily slid into a designed gap and could be reused.

4 Results and discussion

4.1 Design and procedure for a magnetophoresis‑based platform under a static‑fluid environment

A magnetophoresis-based microfluidic platform under a static-fluid environment consists of a permanent magnet, a manually operated syringe, and a microfluidic device which is fabricated with PDMS. Figure 2a shows the design of a microfluidic device for magnetophoretic assays under a static-fluid environment. This device has two paral-lel microchannels, connected with a single syringe through the flexible tubing for the same environment, which are two units for detecting each target sample. While a pair of microchannels share one end which is connected with a manually operated syringe, opposite ends of the micro-channels are designed to be open chambers for easy injec-tion of sample containing paramagnetic solution and micro-beads. The width, length, and height of the microchannels were 1 mm, 6 mm, and 100 μm, respectively. For the same environment, the whole length of two microchannels was designed to be identical. The diameter of the open cham-ber for sample injection was 3 mm. A nickel microstruc-ture combined with a permanent magnet, which serves to enhance a magnetic field gradient acting on microbeads, was arranged in the perpendicular direction to the micro-channels (Fig. 2b). Micropillar structures at the initial posi-tion near the nickel microstructure were used to align the injected microbeads (Fig. 2c). The spacing between each micropillar structure was 10 μm, which is sufficient to line up the microbeads (with a diameter of 20 μm) dragged by manual operation of a syringe. In the fabrication process, a large aspect ratio of the micropillar structure causes a bend-ing of the structure. The bent micropillar structures could not align the microbeads due to a large interval. Therefore, the height of the microchannel was designed to be 100 μm in order to prevent the micropillar structure from bending.

The injected microbeads sank down to the bot-tom in the microchannel. For minimizing the interac-tion between the surface of the microchannel and the injected microbeads, the microchannels were treated with Teflon AF. A chemically inert layer of Teflon AF, coated on the microchannel, prevents the adsorption of biotin-coated microbeads (Papautsky et al. 2005). To prevent generation of bubbles due to the hydrophobicity of Teflon-coated microchannel, ethanol was first dropped into four inlets (open chambers) prior to injection of the sample. When the ethanol was filled to the microchan-nel by capillary effect, the ethanol was washed out with a paramagnetic solution. Samples were then dropped into the inlets (Fig. 3a, d). This sample was prepared with MNP-labeled microbeads in the 100 mM Gd-DTPA paramagnetic solution. In this study, Gd-DTPA, which was clinically used as a contrast agent in magnetic reso-nance imaging (MRI), was used as paramagnetic solu-tion. Since Gd-DTPA does not cause protein denatura-tion, it has already been used in various experiments, including protein detection (Hahn and Park 2011) and particle and cell separation (Kang et al. 2008; Shen et al. 2012). The microbeads were first aligned at the initial position, in which micropillar structures were located. This process was performed by applying a negative pressure by gradually pulling the syringe by hand. The drained microbeads were filtered through the micropillar structures which were close to the nickel microstructure (Fig. 3b, e). Figure 3g shows the filtered microbeads by the micropillar structures. When the microbeads were sufficiently aligned with the pillar structures, the syringe operation was stopped, and then a permanent magnet was located near the nickel microstructure (Fig. 3c, f). The microbeads aligned at the initial position moved away from the external magnetic field mainly due to the magnetic repulsive force caused by the difference in magnetic susceptibility between the microbeads and the paramagnetic solution. Here, the magnetic repulsive force is proportional to the volume of the microbeads so that 20-μm beads have been selected to minimize the detection time by 10 min. The magnetic susceptibility of polystyrene microbeads and 100 mM Gd-DTPA solution is −8.75 × 10−6 and 25.96 × 10−6, respectively. The magnetic susceptibility was calculated by the following Eq. (10) (Zhang et al. 2005):

where CGd−DTPA is the concentration of Gd-DTPA (mM). The viscosity of 100 mM Gd-DTPA is about 1.2 × 10−3 Ps·s (Kang et al. 2008). When the microbeads passed through the region 1 mm far from the initial position, the time (t1 mm) was measured to calculate the mean velocity of the microbeads (Fig. 3h).

(10)χm = (0.35CGd−DTPA − 9.04)× 10−6

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4.2 Evaluation on the influence of residual flow

We observed the movement of microbeads with a diameter of 20 μm which were moved away from the initial posi-tion by magnetic repulsive force (Fig. 4a). The moving dis-tance of the microbeads for 9 min was different for each device. As proved in this result, the measurement method based on the moving distance of the microbeads can-not be applied to the manual operation of fluid due to its

variations. We therefore observed mean velocities of the microbeads for the same moving distance from the initial position. The time (t1 mm), for which microbeads move to 1 mm distance, was measured to calculate the mean veloc-ity (Fig. 4b). As expected, the velocity of the microbeads in each device was also different. This difference in velocity for each device is caused by residual flow, which is mainly caused by a small dimension of microchannel and air bub-ble in the device. Manual operation of the syringe causes

Fig. 3 Experimental procedure. a, d Samples including microbeads and paramagnetic solution were dropped into open chambers. b, e Injected microbeads were pull and aligned at the micropillar array by a manually operated syringe. c, f After stopping the syringe, a magnet

was located near microchannels. g Pictures of aligned microbeads. h The velocity of the microbeads was measured by observing the trave-ling time for a distance of 1 mm

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a rapid change in pressure (e.g., negative pressure, ∆P) at one end of the fluid channel within the microfluidic device, which induces the flow in the microchannel to release the negative pressure. However, due to the small dimension of the microchannel, the high hydraulic resistance (Rh) regu-lates the fluid flow (∆P = RhQ), which increases the time required to stabilize the pressure difference. This hydraulic resistance is generally inversely proportional to the width and the third powers of the height and is proportional to the length of the microchannel. Furthermore, a rapid pressure change during manual operation expands the compressive materials, which cause an increase in the size of the small air bubbles present. As the pressure in the microchannel stabilizes, air bubbles return to their original size, which in turn causes a residual flow. In addition, the residual flow could be affected by various uncontrollable parameters such as Tygon tube, PDMS chip, and syringes. For inves-tigating the residual flow, rhodamine B-marked microbe-ads with a diameter of 2 μm were mixed in the paramag-netic solution and the movement of the microbeads was observed for 20 min after stopping the manual operation

of the syringe (Fig. 4c). As a result, the residual flow was sharply decreased after stopping the manual operation, but still remained until 20 min later (Fig. 4d). In addition, the residual flow was different for each device because an ini-tial pressure change by manual operation also affects the residual flow. For an ideal static-fluid environment without residual flow affecting the results, a stabilization time of more than at least 20 min is required. Therefore, a method without stabilization time is required for rapid detection.

4.3 Compensation with a reference microchannel

The residual flow should be methodologically eliminated by comparison with negative control in the same envi-ronment. Therefore, we first simulated a magnetic field profile at the designed four channels (sample and refer-ence channels of units 1 and 2) using the Finite Element Method Magnetics (FEMM) program to check whether the microchannels have the same magnetic environment (Fig. 5a). Figure 5b shows the profile of the magnetic flux density along the x-axis from the initial position at each

Fig. 4 Influence of residual flow. a Moving distance of 20-μm beads from the initial position by magnetic force. b Velocity of 20-μm beads for the same moving distance (1 mm). c Observation

of the residual flow using 2-μm beads (rhodamine B-marked). Scale bar = 100 μm. d The velocity of 2-μm beads for 20 min after stop-ping of manual operation

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microchannel. As a result, we could know that the designed four microchannels have a similar magnetic profile, which presents the microchannels with the same magnetic field environment. In addition, two microchannels for each unit, the sample channel for the assay using MNP-labeled microbeads and the reference channel for a reference with bare microbeads, were designed to have the same hydrody-namic environment by connecting the only single syringe. Then, we confirmed that both microchannels have the same environment by injecting the same bare microbeads into the two microchannels. Figure 5c shows well-aligned microbe-ads at the initial position under no magnet. These microbe-ads in both microchannels had remained aligned state at the initial position over time (Fig. 5d). On the other hand, when a magnet was placed near the microchannels, the microbe-ads in both microchannels moved away from the magnet at

the same velocity (Fig. 5e, f). These results show that two microchannels had the same environment and also repre-sent that the residual flow in both microchannels was the same, and thus, the residual flow could be considered as an offset signal. Accordingly, the variation of velocity caused by the residual flow could be eliminated by the velocity dif-ference between the microchannels for the reference and the sample.

4.4 Comparison of the velocity of the microbeads to eliminate the influence of residual flow

To demonstrate the detection of labeled MNPs by com-paring the velocity between microchannels as a reference and sample, MNP-labeled microbeads were prepared by binding between streptavidin-functionalized MNPs and

Fig. 5 a Simulation result using the Finite Element Method Magnetics (FEMM). b The pro-file of the magnetic flux density in each microchannel. c–f A pair of microchannels under the same environment to remove the influence of the residual flow c, d under no magnet e, f under a magnet

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biotin-functionalized microbeads. Eighty μL of biotin-coated microbeads (with a diameter of 20 μm and a con-centration of 3.08 × 106 beads/mL) was incubated with 0.45 mg/mL of streptavidin-coated MNPs (with a diameter of 130 nm) for 1 h. After the reaction, the solution was fil-tered with a 0.2 μm nylon filter (Corning Inc., Corning, NY, USA) to eliminate the unbounded free MNPs. Then, bare microbeads were injected into the reference micro-channel, while MNP-labeled microbeads were injected into the sample microchannel. When the microbeads in each microchannel were aligned at an initial position near the nickel microstructure, the manual operation of the syringe was stopped and a permanent magnet was placed adjacent to the nickel microstructure. The 1 mm distance traveling time of each microbead was observed to estimate the veloc-ities of microbeads, and the observation time was within 5 min under all experimental conditions. To reduce the chance of contact between the microbeads, we controlled the concentration of microbeads to inject approximately 30 microbeads into each microchannel. Nevertheless, two or more aggregated microbeads were found, but they did not show any significant difference in velocity; hence, they were included in the data.

Figure 6a shows the results of repeated experiments of five devices. The velocity of the bare microbeads mov-ing away from the initial position only by the magnetic repulsive force in the reference microchannel was faster than that of MNP-labeled microbeads in the sample

microchannel due to the countercurrent magnetic attractive force acting on the labeled MNPs. Compared to the veloc-ity of MNP-labeled microbeads in each device, the differ-ent velocities from five microfluidic devices were obtained despite the use of the same samples. However, it was con-firmed that the variation in the velocities of MNP-labeled microbeads was similar to the variation in the velocity of bare microbeads in the reference microchannel. As a result, the effect of residual flow could be considered as an off-set signal. Therefore, the variation caused by residual flow could be calibrated using the velocity difference of the microbeads between reference and sample microchannels. Figure 6b shows the velocity difference of microbeads for each device from Fig. 6a data. The velocity difference was calculated by vbare − vMNP and the standard deviation was

calculated by √

s2bare

n+

s2MNP

m, where vbare and vMNP are the

mean velocity of bare beads and MNP-labeled beads, and sbare and sMNP are the standard deviation of the velocity of bare beads and MNP-labeled beads, and n and m are the number of injected bare beads and MNP-labeled beads, respectively. While the coefficient of variation (CV) of the mean velocities of the MNP-labeled microbeads for five microfluidic devices was 26.6%, the CV of the average val-ues of the velocity differences of the microbeads was con-siderably reduced to 5.9%.

Next, we investigated the detection of five different concentrations of labeled MNPs. For the preparation of

Fig. 6 a, b Experimental results for a MNP concentration of 0.45 mg/mL. a A velocity of bare microbeads in the refer-ence channel and MNP-labeled microbeads in the sample chan-nel for five microfluidic devices. Comparing the velocity of the microbeads in each device, the velocity was different despite the use of the same samples (n ≥ 20). b The velocity differ-ence between bare microbeads and MNP-labeled microbeads was similar in five devices. The velocity difference means vbare − vMNP. c Mean velocities of MNP-labeled microbeads in five different MNP concentra-tions (n = 4). d Average values of the velocity differences in five different MNP concentra-tions (n = 4)

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microbeads conjugated with the respective concentra-tion of MNPs, 80 μL of biotin-coated microbeads (with a diameter of 20 μm and a concentration of 3.08 × 106 beads/mL) was incubated with 0.06, 0.12 , 0.24, 0.45, and 0.83 mg/mL of streptavidin-coated MNPs (with a diameter of 130 nm) for 1 h. The procedures were identical to one of the previous experiments. Figure 6c shows the mean velocities of MNP-labeled microbeads in five different MNP concentrations. In the case of 0.83 mg/mL of MNPs, the MNP-labeled microbeads did not move away and stayed at the initial position because the magnetic attrac-tive force acting on labeled MNPs was bigger than the magnetic repulsive force acting on a microbead. Accord-ingly, the mean velocity of the samples which stayed at the initial position was marked as zero. Meanwhile, MNP-labeled microbeads in other concentrations showed the movement, and the concentrations of the labeled MNPs could not be classified with only the mean velocity of MNP-labeled microbeads. However, the difference in mean velocity between bare microbeads and samples can classify the samples by the concentration of the labeled MNPs. Figure 6d shows the average values of velocity dif-ferences in five different MNP concentrations. The error bars are the standard deviation of the velocity differences of four devices. As a result, in the case of 0.06 mg/mL of MNPs, the average value of the velocity difference turned out to be nearly zero because the magnetic attractive force was not sufficient to affect the velocity by small quanti-ties of labeled MNPs. Meanwhile, in the case of 0.83 mg/mL of MNPs, the average value of velocity difference was the same as the mean velocity of bare microbeads because MNP-labeled microbeads showed no movement (the velocity was zero) by too many labeled MNPs. In the case of 0.12, 0.24, and 0.45 mg/mL of MNPs, while the respec-tive velocities of the microbeads in each sample did not show a tendency in the results due to the errors by residual flow (Fig. 6c), their average values of velocity differences were presented for each concentration with distinguish-able results (Fig. 6d). In addition, the standard deviation was dramatically reduced by comparing the velocities of the microbeads between the microchannels for the refer-ence and the sample. Consequently, this magnetophoresis-based microfluidic platform has been shown to be used for quantitative detection without a dynamic flow environment which requires delicate control of electrical pumps.

5 Conclusions

We demonstrated a magnetophoresis-based microfluidic platform under a static-fluid environment, which can be operated without an external power source such as electri-cal pumps and be used to detect the concentration of target

samples based on MNPs. Contrary to the conventional magnetophoresis using a dynamic flow environment which requires delicate control of pumps, this platform only worked with a simple manual operation using syringes. To calibrate the variation due to the uncontrollable residual flow caused by manual operation, this platform has a pair of microchannels; a reference microchannel and a sample microchannel. The variation could be eliminated by com-paring the velocities between bare microbeads in the ref-erence microchannel and MNP-labeled microbeads in the sample microchannel. Consequently, we could classify five different concentrations of MNP-labeled microbeads. Based on these results, this magnetophoresis-based micro-fluidic platform under a static-fluid environment would be developed to have a higher resolution and sensitivity and could be exploited for a portable detection platform. Addi-tionally, we expect the presented device to be extended to develop a more practical device for discriminating com-plex target molecules such as MNP protein microbead complexes.

Acknowledgements This research was supported by a Mid-Career Researcher Program (Grant No. NRF-2016R1A2B3015986) and a Bio and Medical Technology Development Program (Grant No. NRF-2015M3A9B3028685) through the National Research Foundation of Korea funded by the Ministry of Science, ICT and Future Planning. The authors also acknowledge a KAIST Systems Healthcare Program.

References

Adams JD, Kim U, Soh HT (2008) Multitarget magnetic activated cell sorter. Proc Natl Acad Sci USA 105:18165–18170. doi:10.1073/pnas.0809795105

Chon H, Lim C, Ha S-M, Ahn Y, Lee EK, Chang S-I, Seong GH, Choo J (2010) On-chip immunoassay using surface-enhanced Raman scattering of hollow gold nanospheres. Anal Chem 82:5290–5295. doi:10.1021/ac100736t

Emaminejad S, Dutton RW, Davis RW, Javanmard M (2014) Mul-tiplexed actuation using ultra dielectrophoresis for proteom-ics applications: a comprehensive electrical and electrothermal design methodology. Lab Chip 14:2105–2114. doi:10.1039/c4lc00036f

Gao Y, Lam AWY, Chan WCW (2013) Automating quantum dot bar-code assays using microfluidics and magnetism for the devel-opment of a point-of-care device. ACS Appl Mater Interfaces 5:2853–2860. doi:10.1021/am302633h

Godino N, Snakenborg D, Kutter JP, Emnéus J, Hansen MF, Muñoz FX, Javier del Campo F (2009) Construction and characterisa-tion of a modular microfluidic system: coupling magnetic cap-ture and electrochemical detection. Microfluid Nanofluid 8:393–402. doi:10.1007/s10404-009-0468-8

Hahn YK, Park J-K (2011) Versatile immunoassays based on isomag-netophoresis. Lab Chip 11:2045–2048. doi:10.1039/c0lc00569j

Hahn YK et al (2007) Magnetophoretic immunoassay of allergen-specific IgE in an enhanced magnetic field gradient. Anal Chem 79:2214–2220. doi:10.1021/ac061522l

Hahn YK, Chang J-B, Jin Z, Kim H-S, Park J-K (2009) Magneto-phoretic position detection for multiplexed immunoassay using

Microfluid Nanofluid (2017) 21:74

1 3

Page 11 of 11 74

colored microspheres in a microchannel. Biosens Bioelectron 24:1870–1876. doi:10.1016/j.bios.2008.09.016

Han D, Hwang H, Park J-K (2013) Optoelectrofluidic behavior of metal–polymer hybrid colloidal particles. Appl Phys Lett 102:054105. doi:10.1063/1.4790622

Ino K et al (2008) Cell culture arrays using magnetic force-based cell patterning for dynamic single cell analysis. Lab Chip 8:134–142. doi:10.1039/b712330b

Jin Z et al (2009) Magnetic nanoclusters for ultrasensitive mag-netophoretic assays. Small 5:2243–2246. doi:10.1002/smll.200900311

Jo Y, Shen F, Hahn YK, Park J-H, Park J-K (2016) Magnetophoretic sorting of single cell-containing microdroplets. Micromachines 7:56. doi:10.3390/mi7040056

Kang JH, Choi S, Lee W, Park J-K (2008) Isomagnetophoresis to discriminate subtle difference in magnetic susceptibility. J Am Chem Soc 130:396–397. doi:10.1021/ja0770678

Lee H, Kim J, Kim H, Kim J, Kwon S (2010) Colour-barcoded mag-netic microparticles for multiplexed bioassays. Nat Mater 9:745–749. doi:10.1038/nmat2815

Lin Y-H, Peng P-Y (2015) Semiconductor sensor embedded micro-fluidic chip for protein biomarker detection using a bead-based immunoassay combined with deoxyribonucleic acid strand labe-ling. Anal Chim Acta 869:34–42. doi:10.1016/j.aca.2015.03.002

Lund-Olesen T, Dufva M, Hansen MF (2007) Capture of DNA in microfluidic channel using magnetic beads: increasing capture efficiency with integrated microfluidic mixer. J Magn Magn Mater 311:396–400. doi:10.1016/j.jmmm.2006.10.1171

Papautsky I et al (2005) Adsorption of fluorescently labeled microbeads on PDMS surfaces. Proc SPIE 5718:159. doi:10.1117/12.591781

Rissin DM et al (2010) Single-molecule enzyme-linked immunosorb-ent assay detects serum proteins at subfemtomolar concentra-tions. Nat Biotechnol 28:595–599. doi:10.1038/nbt.1641

Robert D, Pamme N, Conjeaud H, Gazeau F, Iles A, Wilhelm C (2011) Cell sorting by endocytotic capacity in a microfluidic

magnetophoresis device. Lab Chip 11:1902–1910. doi:10.1039/c0lc00656d

Saliba A-E et al. (2010) Microfluidic sorting and multimodal typing of cancer cells in self-assembled magnetic arrays. Proc Natl Acad Sci USA 107:14524–14529. doi:10.1073/pnas.1001515107

Sasso LA, Johnston IH, Zheng M, Gupte RK, Undar A, Zahn JD (2012) Automated microfluidic processing platform for mul-tiplexed magnetic bead immunoassays. Microfluid Nanofluid 13:603–612. doi:10.1007/s10404-012-0980-0

Shen F, Hwang H, Hahn YK, Park J-K (2012) Label-free cell sepa-ration using a tunable magnetophoretic repulsion force. Anal Chem 84:3075–3081. doi:10.1021/ac201505j

Tang J, Tang D, Niessner R, Chen G, Knopp D (2011) Magneto-controlled graphene immunosensing platform for simultaneous multiplexed electrochemical immunoassay using distinguishable signal tags. Anal Chem 83:5407–5414. doi:10.1021/ac200969w

Tenje M et al (2015) Acoustic trapping as a generic non-contact incu-bation site for multiplex bead-based assays. Anal Chim Acta 853:682–688. doi:10.1016/j.aca.2014.07.008

Tennico YH, Hutanu D, Koesdjojo MT, Bartel CM, Remcho VT (2010) On-chip aptamer-based sandwich assay for thrombin detection employing magnetic beads and quantum dots. Anal Chem 82:5591–5597. doi:10.1021/ac101269u

Zhang H, Moore LR, Zborowcki M, Williams PS, Margel S, Chalm-ers JJ (2005) Establishment and implications of a characteriza-tion method for magnetic nanoparticle using cell tracking veloci-metry and magnetic susceptibility modified solutions. Analyst 130:514–527. doi:10.1039/b412723d

Zhang DS-z et al (2016) Dual-encoded microbeads through a host-guest structure: enormous, flexible, and accurate barcodes for multiplexed assays. Adv Funct Mater 26:6146–6157. doi:10.1002/adfm.201601963

Zhu T, Marrero F, Mao L (2010) Continuous separation of non-mag-netic particles inside ferrofluids. Microfluid Nanofluid 9:1003–1009. doi:10.1007/s10404-010-0616-1