Handheld electrochemical sensors are part of the daily routine for millions of people with diabetes around the globe who monitor their blood sugar levels with electric glucometers.
While such sensors have revolutionized at-home medical testing for diabetics, they have not yet been successfully applied to diagnosing other conditions.
Sensors like glucometers detect glucose in blood based on the activity of an enzyme, and there are only a limited number of enzymes that can be used to sense biomarkers of human disease.
An alternative detection strategy based on binding events between antibodies and their molecular targets have been investigated to expand the use of electrochemical sensors for medicine, but these sensors fall victim to the rapid accumulation of “fouling” substances from biological fluids on their conductive surfaces, which deactivate them.
Existing antifouling coatings are difficult to mass-manufacture, suffer from quality and consistency issues, and are not very effective.
Now, a new diagnostic platform technology developed by researchers at the Wyss Institute for Biologically Inspired Engineering at Harvard University known as “eRapid” enables the creation of low-cost, handheld electrochemical devices that can simultaneously detect a broad range of biomarkers with high sensitivity and selectivity in complex biological fluids, using as little as a single drop of blood. The technology is described in the newest issue of Nature Nanotechnology.
“As long as an antibody exists for a given target molecule, eRapid can detect it,” said co-author Pawan Jolly, Ph.D., a Senior Research Scientist at the Wyss Institute. “By solving the biofouling problem with a simple yet robust design, we are now able to easily mass-produce biochemical sensors for a wide variety of applications at low-cost.”
The challenge in developing the antifouling coating was to prevent accumulation of off-target substances on the sensor’s metal electrodes while still maintaining their conductivity to allow sensing of the target.
After experimenting with a variety of recipes, the research team developed a simple, porous, 3D matrix consisting of bovine serum albumin (BSA) crosslinked with glutaraldehyde and supported by a network of conducting nanomaterials, such as gold nanowires or carbon nanotubes.
The small pore size of the BSA matrix size-excludes proteins found in blood and plasma, and the BSA’s weak negative charge prevents the strong adhesion of positively charged biomolecules onto the sensor.
When the researchers tested their nanomaterial-coated sensors in human blood serum and plasma, they retained more than 90% of their ability to detect signal even after being stored for one month in those biofluids, whereas sensors coated with best previously published anti-fouling coatings lost significant signal sensitivity when incubated for one hour, and were completely inactivated after one day.
To functionalize the coated sensors, the researchers attached antibodies to the surface of the nanomaterial coating on top of the electrode, and used a “sandwich assay” to convert the antibody binding event into a chemical signal that precipitates onto the electrode surface, thereby generating an electric signal.
The magnitude of the electrical signal directly correlates to the amount of the precipitate produced, and thus to the number of target molecules bound to the antibodies, allowing the concentration of the target to be measured.
The team demonstrated the commercial utility of this approach by creating a multiplexed sensor with three separate electrodes, each coated with the BSA/gold nanowire matrix and a layer of antibodies against a specific clinically relevant target molecule: interleukin 6 (IL6), insulin, or glucagon. When they incubated the sensor with the respective target molecules in undiluted human plasma, they observed excellent electrical signals with picogram-per-mL sensitivity.
Conversely, electrodes coated with a published “PEG-SAM” anti-fouling coating failed to produce distinct signals, indicating that they had been irreversibly fouled by off-target molecules in human plasma samples. In addition, the BSA/gold-nanowire-coated sensors can be washed and reused multiple times with minimal signal loss, allowing serial monitoring of biomarkers easily and at low cost.
Since then, the Wyss team has been able to detect more than a dozen different biomarkers ranging from 100 Da to 150,000 Da in size with eRapid, and they are continuing to experiment with conductive nanomaterials to optimize the electrode coating and the system’s performance, as well as reduce the cost even further.
They are actively exploring commercialization options for eRapid in the handheld point-of-care diagnostics space, but also hope to extend the coating and sensor technology platform to other targets and contexts, including in-hospital diagnostics, environmental toxin sensing, small molecule detection, and implantable medical devices.
Interestingly, the team – led by the Wyss Institute’s Founding Director Donald Ingber, M.D., Ph.D. – did not originally set out with this goal in mind. This work began because they needed to simultaneously detect multiple biomolecules produced by various types of tissue cells growing within human Organs-on-Chips to non-invasively assess their function and inflammatory status over time. The tiny volume of liquid outflows from the chips’ channels necessitated highly sensitive sensors that could also be multiplexed, which led to the creation of the current technology.
“eRapid emerged from pursuing one innovation that led to another that has the potential to transform medical diagnostics. Hopefully, this simple technology will enable great advances in our ability to develop handheld diagnostic devices that can be used at home, as well as in pharmacies, ambulances, doctor’s offices, and emergency departments in the near future,” said Ingber who is also the Judah Folkman Professor of Vascular Biology at Harvard Medical School and the Vascular Biology Program at Boston Children’s Hospital, and Professor of Bioengineering at Harvard’s John A. Paulson School of Engineering and Applied Sciences.
Additional authors of the paper include former Wyss members Jonathan Sabaté del Río, Ph.D., who is currently a Postdoctoral Fellow at the Institute for Basic Science in Korea and Olivier Henry, Ph.D., who is currently a Program Manager at Imec in Belgium. This research was supported by the Wyss Institute for Biologically Inspired Engineering, the Defense Advanced Research Projects Agency, the Institute for Basic Science, the KeepSmilin4Abbie Foundation, and the National Science Foundation.
eRapid is an electrochemical sensing platform that uses a novel antifouling coating to enable low-cost, multiplexed detection of a wide range of biomolecules for diagnostics and other applications. Credit: Wyss Institute at Harvard University
As a disease attacks a person, physiological signals that represent the biological state of the person change in response to the status of the disease.
A biomarker is a characteristic that is objectively measured and evaluated as an indicator of normal biological processes, pathogenic processes, pharmacologic responses to therapeutic intervention or any measurable diagnostic indicator for assessing the risk or presence of a disease.1 It can include mRNA expression profiles, circulating DNA and tumor cells, proteins, proteomic pattern, lipids, metabolites, imaging methods or electrical signals.2–5
These signals/biomarkers may be obtained from sources such as urine, blood and tissues. Disease biomarker detection that is desired to be accurate, relatively noninvasive and easy to perform, even in point-of-care (POC) settings, can improve the screening, diagnosis, prognosis and recovery on treatment of various diseases.
Acute infectious diseases caused by pathogenic organisms such as bacteria, viruses, fungi and parasites have been a major cause of global death and high disability rates throughout the human history.6, 7 In developing nations, even curable infectious diseases pose a great threat to patients due to lack of affordable diagnosis.8 According to a global report on infectious disease of poverty (2012) by World Health Organization (WHO), each year infectious diseases kill 3.5 million people, mostly the poor and young children who live in low and middle-income countries.9
Over 95% of deaths by infectious disease are due to the lack of proper diagnosis and treatment, and difficulty in accessing adequate healthcare infrastructures.8 Along with infectious diseases, cancer, the uncontrolled growth of abnormal cells which can spread and invade other parts of the body through the blood and lymph system, also figures among a leading cause of death worldwide with 8.2 million deaths in 2012, according to WHO.10
Annual cancer cases are expected to rise from 14 million in 2014 to 22 million within next 2 decades. Similar to infectious diseases, high incidence of cancer occurs in developing nations. According to WHO, 8 million (57%) of new cancer cases, 5.3 million (65%) of cancer deaths and 15.6 million (48%) of 5-year prevalence cancer cases, occurred in less developed regions.11
Infectious diseases and cancer along with other diseases are mostly diagnosed by biomarker detection in laboratories using conventional tests such as enzyme linked immunosorbent assay (ELISA), immunofluorescence, western blotting, immunodiffusion, polymerase chain reaction (PCR), flow cytometry and a wide range of other techniques.12–14 However, most of these assays are complex, take hours to complete, consume large volumes of samples and reagents, and require bulky and expensive instruments limiting their applications in rural areas and developing nations.
Therefore, simple, low-cost, portable diagnostic devices and methods, especially point-of-care (POC) diagnostic devices that offer great potential to detect and monitor diseases, even at resource-limited settings are in great need.
Development of POC devices for simple, timely and early disease diagnosis can prevent the spread of infectious diseases, and decrease cancer fatality, as many cancer patients (including breast, colorectal, oral and cervical) have high chance to be cured if detected early and treated adequately.
WHO has developed a list of general characteristics that make a diagnostic test appropriate for resource-limited sites, abbreviated as ASSURED, and includes, Affordable by those at risk, Sensitive, Specific, User-friendly, Rapid treatment and robust use, Equipment-free and finally Delivered to those who need it.15
Microfluidics technology possesses remarkable features for simple, low-cost, and rapid disease diagnosis, such as low volumes of reagent consumption, fast analysis, high portability along with integrated processing and analysis of complex biological fluids with high sensitivity for health care application.16–22
An enormous number of microfluidic devices have been developed for biomedical applications.23–29 These devices enable on-chip POC diagnosis and real-time monitoring of diseases from a small volume of body fluids.
These microfluidic devices may act as a bridge to improve the global health care system with high efficiency and sensitivity, especially for remote areas with low-resource settings, such as the underdeveloped and developing countries, in home health care setting, and in emergency situations.
Because of all these significant features, numerous microfluidic devices have been developed for the biomarker detection in disease diagnosis, which includes different types of cancer30–32 from colorectal carcinoma33, 34 and hepatocellular carcinoma32 to ovarian cancer33, 35 and prostate cancer,36, 37 different types of infectious diseases from food-borne pathogen38 and Hepatitis B39 to meningitis40, 41 and dengue virus,42 and other diseases from cardiovascular disease43, 44 to Alzheimer’s diseases. 45.
These microfluidic platforms includes glass,21, 46, 47 polydimethylsiloxane (PDMS),45, 48, 49 poly(methyl methacrylate) (PMMA),36, 50, 51 poly(cyclic olefin),52, 53 paper-based,23, 54–59 and hybrid devices.36, 60, 61
This article reviews recent advances of biomarker detection for disease diagnosis using microfluidic techniques. It first introduces different microfluidic POC platforms used for disease biomarker detection with a brief introduction of their common fabrication techniques. Because of their ease of fabrication, cost-effective characteristics, and broad applications in disease diagnosis, it mainly focuses on cost-effective microfluidic platforms such as polymer (e.g. PDMS and PMMA) and paper-based microfluidic platforms.
Next, it highlights various detection strategies for disease biomarker detection using microfluidic devices, including colorimetric, fluorescence, chemiluminescence, electrochemiluminescence (ECL), and electrochemical detection. Lastly, we briefly discuss the future trends of this field. Although microfluidic platforms have great potential for the diagnosis of a broad range of diseases, this article emphasizes the applications of microfluidic devices in infectious diseases and cancer.
Microfluidic platforms for biomarker detection
In the early stage of microfluidics, microfluidic devices were predominantly made with methods borrowed from microelectronics field and involved materials such as glass, quartz or silicon. Silicon and glass are more expensive and less flexible to work with, as compared to polymers (e.g. PMMA and PDMS). Most of them have good optical properties similar to glass, but their fabrication (e.g. soft lithography62–64) does not have stringent requirements on cleanroom facility, which makes polymer-based microfluidic devices widely used. Within recent years, paper-based microfluidic devices have debuted as a lower-cost microfluidic platform.19, 65, 66
The choice of material depends on the research application, detection system, fabrication facility, cost and other factors such as resistance to different chemicals, thermal conductivity, dielectric strength and sealing properties. This section mainly aims to give a general introduction of various cost-effective microfluidic platforms used for disease biomarker detection. Since the focus of this article is not to review recent fabrication techniques, only common fabrication techniques and their recent advances are briefly described. A few other review articles described more details of fabrication methods for different microfluidic platforms.67–69
PDMS Microfluidic Platforms
PDMS is one of the most widely used elastomers for microfluidic devices as it is optically transparent, elastic, and cures at low temperature. It can seal with itself and a range of other materials after being exposed to air plasma. The ease and low cost of fabrication and ability to be cast in high resolution add to its advantages. In contrast to other thermoplastic materials, PDMS is gas permeable, making it compatible for cell culture. Although PDMS is one of the most widely used cost-effective microfluidic platforms, there are some limitations of PDMS as well. PDMS swells in organic solvents and low molecular weight organic solutes. It cannot withstand high temperature and the mechanical resistance is quite low. There are different methods available for the fabrication of PDMS devices including soft lithography, casting, injection molding, imprinting, hot embossing, laser ablation and others.22, 62, 63
Soft lithography is the most widely used method for PDMS fabrication. Soft lithography can start with creation of a photomask in a transparency film. The resolution of transparency is >20 μm as compared to chrome mask ~500 nm.71
Photoresist is then added to the silicon wafer, and exposed to UV light through the photomask to produce a positive relief of photoresist on a silicon wafer (master).
Masters can also be fabricated by techniques like etching in silicon and electroforming metal. Channels in PDMS can be formed by replica molding once a master is fabricated.
The cured PDMS replica can be bonded with another flat layer of PDMS, glass or other materials to form a closed system. Based on soft lithography, Kung et al.70 demonstrated a novel method for fabricating 3D high aspect ratio PDMS microfluidic networks with a hybrid stamp. Figure 1 shows the schematic of fabrication process flow. A SU8 master is treated with trichloro (1H,1H,2H,2H-perfluorooctyl)silane (PFOCTS) to facilitate subsequent demolding. An uncured PDMS mixture is then poured on the master followed by pressing against the hybrid stamp.
Then, the casted PDMS film is peeled off from the master since it tends to adhere to the hybrid stamp, as there is less PFOCTS on the hybrid stamp. Afterwards, the PDMS film is transferred and bonded with glass/silicon by oxygen plasma treatment, followed with the removal of the supporting PDMS, the polystyrene plastic plate and residual PDMS. Finally, the stacking process is repeated to complete the 3D fabrication.
They showed that multilayer 3D PDMS structures could be constructed and bonded between two hard substrates. As an example, they fabricated a microfluidic 3D deformable channel by sandwiching two PDMS membranes (20 μm wide and 80 μm tall) between two glass substrates. This 3D fabrication method could be applied in electrokinetics, optofluidics, inertial microfluidics, and other fields where the shape of the channel cross-section is significant in device physics. Comina et al.72 described another method for fabrication of 3D PDMS devices using templates printed with a commercial micro-stereo lithography 3D printer with a resolution of 50 μm.
The process eliminates the need for clean room facilities and repeated photolithographic steps required for templates with different thickness. They reported that the templates are reusable and can be fabricated within 20 min, with an average cost of 0.48 US$.
Thermoplastic microfluidic platforms
Thermoplastics are also being used as the substitute of glass and silicon as the microfluidic platform due to their chemical and mechanical properties. Thermoplastic devices are economical for mass production and are compatible with most chemical reagents and biological assays. Several kinds of thermoplastic have been used such as PMMA, (i.e. acrylic), polycarbonate, polyester and polyvinylchloride (PVC), because of their low-cost, desirable optical properties and ease of fabrication. They offer better performances than PDMS under mechanical stress.
They don’t require long fabrication and curing time. These thermoplastic devices can be fabricated easily by cutting the pattern using a CO2 laser cutter followed by bonding with adhesive or heat to form 3D devices. Multilayer devices can be completed and become ready for testing in as little as several hours.73
Cassano et al.74 used vacuum bagging for thermal bonding of thermoplastic microfluidic devices. Vacuum bagging completely eliminates time constrains resulting from using solvents, adhesives, or surface treatments. With fabrication technologies including hot embossing or imprinting,75, 76 laser ablation,77 injection molding78 and soft lithography, dimensions of plastic microchannels can be achieved in the range of 15–30 μm.
Recently, simple methods have been developed for rapid prototyping of thermoplastic microfluidic platforms. For example, Roy et al.79 reported a rapid prototyping technique for fabrication of multilayer microfluidic device using styrenic thermoplastic elastomer (TPE).
They established a proof of principle for valving and mixing with three different grades of TPE using an SU-8 master mold. Miserere et al.80 proposed a strategy for the fabrication of flexible thermoplastic microdevices based on lamination process. Low-cost laminator can be used from master fabrication to microchannel sealing. They demonstrated the process using Cyclo-olefin Copolymer (COC).
Rahmanian et al.81 described rapid desktop manufacturing of sealed thermoplastic microchannels. Patterning was achieved by simply drawing the desired microchannel pattern onto the polymer surface using suitable ink as a masking layer, followed by exposure to solvent vapor to yield a desired depth. The channels were then permanently sealed through solvent bonding of the microchannel chip to a mating thermoplastic substrate. Among these various fabrication methods, two of the most widely used fabrication techniques in the field of microfluidic biomarker detection are discussed in brief in this review.
The hot embossing75, 76 or imprinting is an established method to fabricate microchannels in common polymer such as polystyrene (PS), polyethylene terephthalate glycol (PETG), PMMA, PVC, and polycarbonate. Silicon stamps are the more commonly used embossing tools for the fabrication of these polymeric microfluidic devices. A typical hot embossing setup consists of a force frame, which delivers the embossing force via a spindle and a T-bar to the boss or the embossing master. The microstructures are then transferred from the master to the polymer by stamping the master into the polymer by heating above its glass transition temperature (Tg) in vacuum.75 Alternatively, polymer devices can be imprinted at room temperature with elevated pressure. The master structure is pressed into the thermoplastic substrate with a force (e.g. 20–30 kN in case of PMMA or PC) depending on the type and size of the substrate along with the feature to be imprinted.75 Finally, the master and the substrate are isothermally cooled to a temperature just below Tg and then separated. The resulting plastic microchannel dimensions are the exact mirror image of the silicon stamp when devices are hot embossed.
Laser ablation77, 82 is also one of the rapid prototyping methods for microfluidic devices. In this technique, the polymer is exposed to the high intensity laser beam, which evaporates the material at the focal point that is due to photo-degradation or thermal-degradation or the combination of two. Pulsed laser is typically used, so each laser shot will ablate a defined amount of material, depending on the material type and absorption properties, laser intensity, wavelength and number of passes made across the channel. This process leads to the rough surface of the laser-ablated microchannels and have a rippled appearance, which depends upon the absorption of polymer at excimer wavelength. Very high temperature is reached during ablation and particles are ejected from the substrate creating a void, with small particulates on the surface of the substrate material, while other decomposition products become gases (carbon dioxide and carbon monoxide). Laser ablation may be achieved by two ways. Polymer substrate can be exposed to a laser through a mask. A mask is usually made from the material that does not have significant absorption at the laser wavelength used. In the mask-less process, a polymer substrate is placed on a movable stage and either the focused laser beam or the substrate is moved across in x and y direction as defined in the desired pattern.
Paper-based microfluidic platforms
Paper is a thin sheet of material that is generally produced by pressing together cellulosic or nitrocellulose fibers.65 Paper can transport liquids via capillary effect without the assistance of external forces. Fabrication of paper-based devices is simple and does not require the use of clean-room facilities. Paper has good stackability, which allows the formation of 3D structures for complex assays.
The high surface to volume ratio provided by the macroporous structure in paper improves the immobilization of protein and DNA biomarkers, allowing fast detection. Paper-based microfluidics devices can be fabricated both in 2D and 3D for either horizontal or vertical flow.68 Fabrication of the paper-based devices can be subdivided into two categories: (i) construction of hydrophobic barriers, and (ii) two-dimensional cutting.
Constructing hydrophobic barriers
One of the most widely used methods to prepare paper-based analytical devices (μPADs) is to construct hydrophobic barriers in the hydrophilic paper matrix. In this way, reagents and analytes can be made to flow in a certain path preventing mixing and spreading across the surrounding paper surface and achieve multiplexed assays without the issue of cross contamination. Hydrophobic barriers can be created on paper through either a physical deposition83 or a chemical modification method.84
A number of different fabrication methods have been developed to fabricate μPADs, such as fast photolithography,85, 86 wax-based fabrication techniques,83, 87 printing photolithography,88 PDMS printing,89 saline UV/O3 patterning90, flexographic printing,91 and alkenyl ketene dimer (AKD) printing.84 Examples of wax-based fabrication include wax screen-printing,87 wax dipping,92 and wax printing.83 In wax screen-printing,87 solid wax is rubbed through a screen onto paper filters.
The printed wax is then melted into paper so that the melted wax diffuse into paper to form hydrophobic barriers using a hot plate. In wax dipping,92 ironmould is first prepared by the laser cutting technique.
The designed pattern is then developed into paper by transferring the pattern mould (sealed by magnets) into molten wax. Wax printing, in which the designed pattern is directly printed on paper using a solid ink (or wax) printer,93 is considered to be one of the most promising and attractive wax-based methods, due to its low cost and high potential for massive production.
After printing, the wax-printed paper is incubated in an oven so that the melted wax from the paper surface can penetrate into paper to form well-defined microchannels across the whole thickness of the paper-based device owing to the porous structure of the filter paper.
The time required for the patterned wax on paper to penetrate through depends on the temperature used (5 min at 110 0C, 30 s at 130 0C) and the wax-patterned paper is stable when store under 60 0C.83 In 2014, Sameenoi et al.94 reported one-step polymer screen-printing for microfluidic paper-based devices.
In this process, a polystyrene solution that is applied through the screen penetrates through the paper to form a 3D hydrophobic barrier, defining a hydrophilic analysis zone.
The smallest hydrophilic channel and hydrophobic barrier obtained was found to be 670 ± 50 μm and 380 ± 40 μm, respectively. Among these fabrication methods, photolithography and wax printing are widely used. Wax is inexpensive and non-toxic.83, 87, 92 Recently, paper/polymer hybrid devices have been developed (Figure 2B), but their fabrication methods is mainly derived from a combination of paper-based and polymer microfluidic device fabrication techniques.41, 95
FLASH (Fast Lithographic Activation of Sheets)
One of the most widely used fabrication technology for constructing hydrophobic barriers in paper-based devices is photolithography or FLASH.86 Chromatography paper is the commonly used substrate. FLASH requires a UV lamp, a printer and a hotplate along with photoresist such as SU-8 and other organic solvents. Figure 2A shows the procedures. In this technique86 photoresist is first poured onto a piece of paper and spread evenly and baked on a hotplate at 130 0C for 5–10 min to evaporate propylene glycol monomethyl ether acetate (PGMEA ) from the photoresist.
Then, the paper is covered with a photomask and exposed to UV light. After incubation in an oven, the chromatography paper is developed in acetone, followed with rinsing with isopropyl alcohol. After drying, the paper-based device is ready to use.
2.3.2 Two-dimensional cutting
Another way to create paper-based microfluidic device is 2D cutting. Paper channels are cut through computer controlled X–Y knife plotters or CO2 laser cutters, and then fixed to suitable plastic cassettes to form hybrid devices.96,41 Nitrocellulose, conventional photocopy paper and chromatography paper can be used. Thuo et al.97 described the use of embossing and a “cut-and-stack” method to develop microfluidic devices from omniphobic paper. They demonstrated that fluid flow in these devices was similar to open-channel microfluidic devices and cut layer generated 3D systems.