A microfluidic biosensor platform was designed and fabricated using low-cost dry film photoresist technology for the rapid and sensitive quantification of various analytes. This single-use system allows for the electrochemical readout of on-chip-immobilized enzyme-linked assays by means of the stop-flow technique.
In recent years, biomarker diagnostics became an indispensable tool for the diagnosis of human disease, especially for the point-of-care diagnostics. An easy-to-use and low-cost sensor platform is highly desired to measure various types of analytes (e.g., biomarkers, hormones, and drugs) quantitatively and specifically. For this reason, dry film photoresist technology – enabling cheap, facile, and high-throughput fabrication – was used to manufacture the microfluidic biosensor presented here. Depending on the bioassay used afterwards, the versatile platform is capable of detecting various types of biomolecules. For the fabrication of the device, platinum electrodes are structured on a flexible polyimide (PI) foil in the only clean-room process step. The PI foil serves as a substrate for the electrodes, which are insulated with an epoxy-based photoresist. The microfluidic channel is subsequently generated by the development and lamination of dry film photoresist (DFR) foils onto the PI wafer. By using a hydrophobic stopping barrier in the channel, the channel is separated into two specific areas: an immobilization section for the enzyme-linked assay and an electrochemical measurement cell for the amperometric signal readout.
The on-chip bioassay immobilization is performed by the adsorption of the biomolecules to the channel surface. The glucose oxidase enzyme is used as a transducer for electrochemical signal generation. In the presence of the substrate, glucose, hydrogen peroxide is produced, which is detected at the platinum working electrode. The stop-flow technique is applied to obtain signal amplification along with rapid detection. Different biomolecules can quantitatively be measured by means of the introduced microfluidic system, giving an indication of different types of diseases, or, in regard to therapeutic drug monitoring, facilitating a personalized therapy.
Over the past two decades, diagnostic applications have become elementary for in-depth studies on the development of global public health. Traditionally, laboratory diagnostic tools are used for the detection of diseases. Even though they still play a key role in the diagnosis of diseases, point-of-care testing (POCT) performed near the patient or by the patient himself has become more and more commonplace in recent years. Especially in such cases that require immediate treatment, such as acute myocardial infarction or diabetes monitoring, the rapid confirmation of a clinical finding is essential. Hence, there is a growing need for POCT devices that can be operated by non-experts and that are concurrently capable of performing precise in vitro diagnostic tests in a short time1,2,3,4.
Remarkable improvements have already been achieved in the field of POCT. However, there are still many challenges to overcome5,6,7,8. For a POCT platform to be successfully launched to the market and to be competitive with laboratory diagnostics, the device must strictly fulfill the following requirements: (i) provide precise and quantitative test results that are consistent with laboratory findings; (ii) have short sample-to-result times, enabling the immediate treatment of the patient; (iii) feature uncomplicated and easy handling, even when operated by untrained individuals, and require minimized user intervention; and (iv) comprise of a low-cost sensor unit designed for single-use applications. Furthermore, equipment-free diagnostics are favorable, mainly in resource-poor environments3,4,6.
Due to these severe requirements, only two POCT systems based on electrochemical detection (e.g., blood glucose test strips) and on lateral flow immunoassays (e.g., home pregnancy tests) have been successfully launched to the market so far. However, both systems suffer from disadvantages such as poor performance (i.e., blood glucose monitoring has inaccurate test results and lateral flow assays only provide qualitative (positive or negative) measurement results)4,6. These drawbacks of conventional POCT systems have led to an increasing demand on exploring new technologies that offer fast, low-cost, and quantitative detection at the point of care4,5.
To meet these challenges facing POCT devices, DFR technology has been recently employed for the fabrication of disposable and low-cost biosensors9,10,11,12,13,14. Compared to soft and liquid lithographic materials, such as PDMS or SU-8, DFRs present many benefits: they (i) are available in a variety of compositions and thicknesses (from a few microns to several millimeters); (ii) have a very rough surface area, which facilitates adhesion to various materials; (iii) feature excellent thickness uniformity; (iv) offer cheap, facile, and high-throughput fabrication for mass production; (v) are easy to cut with various low-cost tools, like a simple pair of scissors; and (vi) allow for the creation of three-dimensional structures, such as microfluidic channels, by stacking multiple DFR layers on top of each other.
On the other hand, DFRs in general have a relatively poor resolution compared to liquid photoresists, which is mainly caused by the film thickness and by the increased distance between the mask and the DFR due to the protective foil, which additionally enables light scattering. Still, for the manufacturing of integrated microfluidic biosensors, DFRs are highly suitable for low-cost mass production.
Therefore, we present in this work the fabrication and application of a DFR-based electrochemical microfluidic biosensor. The detailed protocol describes each production step of the biosensor platform, the on-chip immobilization of a DNA-based model assay, and its electrochemical readout using the stop-flow technique. This universal platform enables the detection of numerous kinds of biomolecules, using different assay technologies (e.g., genomics, cellomics, and proteomics) or assay formats (e.g., competitive, sandwich, or direct). Based on such a DFR platform, our group previously successfully demonstrated the rapid and sensitive quantification of various analytes, including antibiotics13,15,16 (tetracycline, pristinamycin, and ß-lactam antibiotics), troponin I17, and substance P18.
1. Fabrication of the Microfluidic Biosensor Using DFR Technology
2. On-chip Assay Immobilization Procedure
3. Amperometric Signal Detection Using the Stop-flow Technique
Design and Fabrication of the Microfluidic Biosensor Platform:
The fabrication of the microfluidic biosensor chips is realized on the wafer-level by standard photolithographic techniques employing multiple DFR layers. This fabrication strategy relies on the lamination of developed layers of DFRs on a platinum-patterned PI substrate, forming the microfluidic channels. A short summary depicting the different fabrication steps is given in Figure 1. A single 6-in wafer comprises 130 microfluidic biosensors, each with a dimension of 8 × 10 mm2.
Figure 1. Graphical illustration of the different fabrication steps of the microfluidic biosensor platform. a) Cut the PI substrate into 6-in round wafers. b) Exposure of a possible spin-coated lift-off resist using the respective mask for platinum patterning. c) Substrate after exposure, post-exposure back, and developing the photoresist. d) Physical vapor deposition of platinum on the substrate. e) Lift-off process to remove the excess photoresist. f) Spin-coating of SU-8, forming an insulation layer. g) O2 plasma process to remove SU-8 residues on the Pt electrodes. h) Galvanic deposition of the Ag/AgCl reference electrode. i) UV exposure and development of different DFR layers. j) Lamination of the DFR layers onto the PI wafer. k) Dispensing solved polytetrafluoroethylene, forming a hydrophobic stopping barrier between the immobilization capillary and the electrochemical cell. l) Final electrochemical microfluidic biosensor. Please click here to view a larger version of this figure.
Each biosensor consists of one microfluidic channel, separated into two distinct areas by a hydrophobic stopping barrier: an immobilization section and an electrochemical cell, marked in red and blue, respectively, in Figure 2. The immobilization part of the microchannel has a surface volume of 10.34 mm2 and a volume of 580 nL, resulting in a high surface-to volume ratio of 155 cm-1. The electrochemical cell includes an on-chip silver/silver chloride reference electrode and a counter and working platinum electrode. This separation of the immobilization area and the electrochemical readout of the assay prevents any contamination of the electrodes with biomolecules and therefore inhibits electrode fouling. Furthermore, it enables the precise metering of the immobilization reagents by capillary filling.
Figure 2. Illustration of the operating principle of the electrochemical microfluidic platform. a) Schematics of a model assay based on avidin. b) Photograph of the microfluidic biosensor showing its main elements, including the counter electrode (CE), the reference electrode (RE), the working electrode (WE), and the stopping barrier (SB). The immobilization area is highlighted in red, and the electrochemical cell is marked in blue. c) Schematic reaction of the oxidation of the produced hydrogen peroxide at the Pt working electrode for amperometric detection inside the electrochemical cell. Please click here to view a larger version of this figure.
By employing DFRs for the fabrication of the sensor, the manufacturing process allows for high-throughput on the wafer-level. Therefore, the costs can be kept down to a minimum. The development of all DFR layers is done using simple and low-cost foil masks and a vacuum exposure unit, rather than costly chrome masks and a mask aligner. In addition, the reduction of the necessary clean-room process steps to a minimum cuts the costs for the fabrication even more. In total, the fabrication procedure takes a workload of roughly 10 h, excluding the baking times and the physical vapor deposition of the platinum.
On-chip Assay Incubation:
The on-chip assay incubation is done only by capillary forces. By pipetting drops of different reagents onto the inlet of the microfluidic channel, the fluids are driven through the channel by capillarity action until they reach the hydrophobic stopping barrier. Under steady-state conditions, the reagents are then incubated for a distinct time. This passive system enables the capillary filling of the channel, without the need of any external instrumentation like a syringe pump, and allows for the precise metering of the reagents, as the fluid automatically stops at the stopping barrier. Also, the workflow is consistent with that of a conventional ELISA and it works with high-viscosity liquids like serum or blood.
For the purpose of this experiment, the chip operation is demonstrated by a simple test assay employing the avidin-biotin interaction, as shown in Figure 2. The on-chip assay incubation starts with the adsorption of the avidin to the capillary surface for 1 h, followed by a blocking step with BSA for another 1 h. Subsequently, 6-FAM/biotin-labeled DNA is incubated in the immobilization capillary for 15 min, where it binds to the avidin molecules. Between each incubation step, any unbound biomolecules are removed by a washing step. The wash buffer is applied via the channel outlet using a custom-made vacuum adapter. In the last step, glucose oxidase-labeled antifluorescein antibodies are introduced to the channel for 15 min. After that, the biosensor is ready for the electrochemical readout, using a 40-mM glucose solution as the substrate.
Amperometric Signal Readout Using the Stop-flow Technique:
For the signal readout of the immobilized assay, the enzymes product (here, H2O2), produced in the presence of its substrate, glucose, can be electrochemically detected at the working electrode in the electrochemical cell. To achieve fast detection along with signal amplification, the so-called stop-flow technique is used13. In this method, the flow of the substrate is stopped, which leads to an accumulation of the product inside the capillary. The amount of produced H2O2 therefore depends on the amount of bound glucose oxidase and is dependent on the amount of bound biotinylated DNA. By restarting the flow, the enhanced H2O2 concentration is subsequently flushed through the electrochemical measurement cell, resulting in a current peak signal, as illustrated in Figure 3.
For data analysis, the stop-flow technique offers two different parameters: the maximum height and the charge (i.e., the integral) of the peak signal. Both parameters are directly proportional to the stopping time and can therefore be used for analyzing the data. Considering the data evaluation, when using the peak height, the signal height depends on the maximum H2O2 concentration and thus on its diffusion coefficient and the applied stop time. The longer the stopping time, the higher the measured signal response. For this reason, the gauged peak height remains constant, down to a minimum length of the immobilization capillary. This special feature of the stop-flow technique allows for a drastic decrease in chip dimensions, particularly in the capillary length, while preserving the sensitivity of the sensor.
Figure 3. Model of a typically obtained electrochemical signal readout of an enzyme-linked assay using the stop-flow technique. a) A flow of 40-mM glucose solution at a rate of 20 µL/min was applied. During the stop phase (1, 2, or 5 min), the enzyme production of H2O2 continues. By restarting the flow, the accumulated H2O2 is flushed through the electrochemical cell, where the hydrogen peroxide is electrochemically detected. By repeating the measurement several times (error bars; SEM), an on-chip calibration curve b) is obtained. Please click here to view a larger version of this figure.
The protocol presented here for the fabrication of a microfluidic electrochemical biosensor enables the development of a low-cost, compact, and easy-to-use platform for the detection of biomolecules. Depending on the assay used afterwards on the biosensor, several different biomarkers can be detected. This makes the platform very versatile and provides broad access to various fields of applications, from standard diagnostic tests (e.g., determining the presence of specific diseases at the doctor's office) to point-of-care applications (e.g., the therapeutic drug monitoring of a patient for individualized drug therapy). Especially in point-of-care diagnostics, miniaturized biosensors have many advantages over conventional methods, which usually require extensive laboratory equipment, specialized employees, and large reagent and sample volumes and have long turnaround times.
The fabrication of this biosensor requires strictly following the protocol; otherwise, manufacturing problems can occur. One of the most often observed issues is the misalignment of the different layers. This can be easily solved by using a more advanced apparatus for the fabrication process, which allows for the easier and more precise alignment of the different layers.
The DFR technology offers the possibility of fast and low-cost fabrication. On the other hand, the technology is limited in terms of resolution. For example, microfluidic channels of very small structures (less than 100 µm) cannot be realized using the protocol presented here. In such a case, the photolithography must be performed by a high-precision mask aligner with glass photomasks. In addition, a too-wide channel can also be problematic, since the channel could bend down, reducing the height of the channel and influencing the microfluidic behavior of the chip.
We believe that DFR technology will be more present in future applications because it can easily be scaled up for commercial use. The mass-production of DFR-based microfluidic sensors will be no hindrance when it hits the market because the technique is well-established in other fields of application (e.g., flexible electronics).
In terms of reducing the complexity of the whole system, our future work will focus on the design and implementation of a disposable microfluidic cartridge that includes the biosensor chip; a waste reservoir; and pre-loaded reagents, such as wash buffer and other assay components. For the amperometric measurement, the delivery of the sample and other reagents can then be provided by pneumatic actuation. This will be performed by the handheld reader and will move through the microfluidic biosensor chip to a waste reservoir.
The authors have nothing to disclose.
The authors would like to thank the German Research Foundation (DFG) for partially funding this work under Grant Numbers UR 70/10-01 and UR 70/12-01.
Material | |||
Pyralux | DuPont | AP8525R | Used as polyimide substrate |
MA-N 1420 | Micro Resist Technology | MA-N1420 | Lift-off resit to define the platinum depostion |
Ma-D 533s | Micro Resist Technology | MaD533S | Developer for MA-N1420 |
Platinum | – | – | Electrode and contact pad material |
Ma-R 404s | Micro Resist Technology | MaR404S | Remover for MA-N1420 |
SU-8 3005 | MicroChem Corp. | SU-8-3005 | Photoresist to define the electrode area and as insulation |
1-methoxy-2-propanol acetate | Sigma-Aldrich | 108-65-6 | Developer for SU-8 3005 |
2-Propanol | VWR | 8.18766.2500 | Removing of the SU-8 developer |
1020R | Ultron Systems Inc. | 1020R | UV sensitive adhesive tape for protection of contact pads |
Arguna S | Degussa | 1935 | For Silver depostion on reference electrode |
KCl | Methrom | 62308.020 | For chloridation of the silver reference electrode |
Pyralux | DuPont | PC1025 | Dry film photoresist |
Sodium carbonat | Fluka | 71352 | Developer for Pyralux PC1025 |
Hydrogen chloride | Sigma-Aldrich | 30720 | To top the development of the DFR |
Teflon AF 1600 | DuPont | AF1600 | For employing the stopping barrier |
Name | Company | Catalog Number | Comments |
Equipment | |||
PA104 | Mega Electronics | – | Bubble etch tank |
FED 53 | Binder | 9010-0018 | Oven |
SPIN150 | APT | – | Spin coater |
Präzitherm | Harry Gestigkeit GmbH | PZ 28-2 | Hot plate |
Hellas | Bungard Elektronik | 40000 | Exposure unit |
Tetra30-LF-PC | Diener | – | Plasma unit |
Univex 500 | Leybold | – | Physical vapor deposition unit |
Shaker S4 | ELMI | – | Orbital shaker |
Sonorex Super 10 P | Bandelin | 783 | Sonic bath |
6221 DC and AC | Keithley | – | Current source |
HRL 350 | Ozatec | – | Laminator unit |
Vaccum pen | EFD | – | Vacuum pen |