What is the difference between using an open or a closed configuration for a microfluidic chip?

What is the difference between using an open or a closed configuration for a microfluidic chip?

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I am doing a college BioMed project on Lungs-on-a-chip to make disease models and I have seen some authors use open configurations, while others use closed configurations. What advantages and disadvantages do they offer on the making of this device?

What is the difference between using an open or a closed configuration for a microfluidic chip? - Biology

Marinke W. van der Helm,‡ a Olivier Y. F. Henry, /> ‡ b Amir Bein, b Tiama Hamkins-Indik, b Michael J. Cronce, /> b William D. Leineweber, /> b Mathieu Odijk, /> a Andries D. van der Meer, /> c Jan C. T. Eijkel, /> a Donald E. Ingber, /> bde Albert van den Berg /> a and Loes I. Segerink /> * a

a BIOS Lab on a Chip group, MESA+ Institute for Nanotechnology, MIRA Institute for Biomedical Technology and Technical Medicine and Max Planck Center for Complex Fluid Dynamics, University of Twente, P. O. Box 217, 7500 AE Enschede, The Netherlands
E-mail: [email protected]

b Wyss Institute for Biologically Inspired Engineering at Harvard University, CLSB Bldg. 5th floor, 3 Blackfan Circle, Boston, USA
E-mail: [email protected]

c Applied Stem Cell Technology, MIRA Institute for Biomedical Technology and Technical Medicine, University of Twente, P. O. Box 217, 7500 AE Enschede, The Netherlands

d Vascular Biology Program and Department of Surgery, Boston Children's Hospital and Harvard Medical School, Boston, USA

e Harvard John A. Paulson School of Engineering and Applied Sciences, Harvard University, Cambridge, USA


Here, we describe methods for combining impedance spectroscopy measurements with electrical simulation to reveal transepithelial barrier function and tissue structure of human intestinal epithelium cultured inside an organ-on-chip microfluidic culture device. When performing impedance spectroscopy measurements, electrical simulation enabled normalization of cell layer resistance of epithelium cultured statically in a gut-on-a-chip, which enabled determination of transepithelial electrical resistance (TEER) values that can be compared across device platforms. During culture under dynamic flow, the formation of intestinal villi was accompanied by characteristic changes in impedance spectra both measured experimentally and verified with simulation, and we demonstrate that changes in cell layer capacitance may serve as measures of villi differentiation. This method for combining impedance spectroscopy with simulation can be adapted to better monitor cell layer characteristics within any organ-on-chip in vitro and to enable direct quantitative TEER comparisons between organ-on-chip platforms which should help to advance research on organ function.

A review on microdroplet generation in microfluidics

Micro-nanofluidic technology is widely used in food safety testing, drug screening, new material synthesis and bioengineering. Droplet microfluidic technology is an important branch of micro-nanofluidic technology. The microdroplet technology can produce high-throughput monodisperse droplets. The droplet technology overcomes many problems of continuous flow, such as small sample reagent volume, no cross-contamination, and rapid chemical reaction. So microdroplets have an important position in the fields of single cell analysis, gene sequencing, and real-time diagnosis. This review focuses on the different methods and applications of microdroplet generation in microfluidics. The droplet generation methods are passive and active methods. The passive method does not require external force, while the active method requires external force such as external electric field, magnetic field, acoustic field, and laser field. It is predicted that the quantitative generation of droplets on demand in microfluidics will be the important direction for future research. This review provides new ideas for the applications of the quantitative generation of microdroplets on demand in microfluidics.

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New automated biological-sample analysis systems to accelerate disease detection

Professor Thomas Gervais of Polytechnique Montréal and his students Pierre-Alexandre Goyette and Étienne Boulais, in partnership with the team led by Professor David Juncker of McGill University, have developed a new microfluidic process aimed at automating protein detection by antibodies. This work, the topic of an article in Nature Communications, points to the arrival of new portable instruments to accelerate the screening process and molecule analysis in biological laboratories to accelerate research in cancer biology.

Microfluidics refers to the manipulation of fluids in microscale devices. Commonly called "labs on a chip," microfluidic systems are used to study and analyze very small-scale chemical or biological samples, replacing the extremely expensive and cumbersome instruments used for traditional biological analyses. Listed in 2001 among the "10 Emerging Technologies That Will Change the World" by the MIT Technology Review, microfluidics is considered just as revolutionary for biology and chemistry as microprocessors have been to electronics and IT, and it applies to a huge market.

Today, this young discipline, which began to take off in the 2000s with closed systems made up of microchannel networks, is itself being radically transformed by the discovery made by the group of researchers from Polytechnique and McGill University, which reinforces the theoretical and experimental foundations of open-space microfluidics.

This technology, which eliminates channels, competes favourably with conventional microfluidics for certain types of analyses. Indeed, the classical configuration of closed-channel microfluidic devices provides several disadvantages: the scale of the channel cross-sections increases the stress that cells undergo when they are culture and they are not compatible with the cell-culture standard, the Petri dish, which makes it hard for the industry to adopt it.

The new approach explored by Polytechnique and McGill University researchers is based on microfluidic multipoles (MFMs), a system of simultaneous fluid suction and aspiration through opposing micro-openings on a very small surface placed in a confined space that is less than 0.1 mm thick. "When they come into contact with one another, these jets of fluid form patterns that can be seen by dyeing them with chemical reagents," says Professor Gervais. "We wanted to understand these patterns while developing a reliable method for modelling MFMs."


To understand these patterns, Professor Gervais's team had to develop a new mathematical model for open multipolar flows. This model is based on a classical branch of mathematics known as conformal mapping that solves a problem related to a complex geometry by reducing it to a simpler geometry (and vice-versa).

PhD student Étienne Boulais first developed a model to study microjet collisions in a multifluidic dipole (an MFM with only two openings), and then, relying on this mathematical theory, extrapolated the model to MFMs with multiple openings. "We can make an analogy with a game of chess in which there is a version with four players, then six or eight, applying a spatial deformation while maintaining the same rules of the game," he explains.

"When subjected to conformal mapping, the patterns created by fluid jet collisions form symmetrical images reminiscent of the paintings of Dutch artist M.C. Escher," adds the young researcher, who has a passion for visual arts. "But far beyond its aesthetic appeal, our model allows us to describe the speed with which molecules move through fluids as well as their concentration. We have defined valid rules for all possible systems configurations of up to 12 poles in order to generate a wide variety of flow and diffusion patterns."

The method is therefore a complete toolbox that will not only make it possible to model and explain the phenomena occurring in MFMs, but also explore new configurations. Thanks to this method, it is now possible to automate open-space microfluidic tests, which up until now have only ever been explored through trial and error.


The design and manufacture of the MFM device was accomplished by Pierre-Alexandre Goyette. This device is a small probe made out of resin using a low-cost 3D printing process and connected to a system of pumps and injectors.

"The expertise of Professor Juncker's team in the detection of proteins by antibodies immobilized on a surface has been invaluable in managing the biological aspects of this project," says the PhD student in biomedical engineering. "The results obtained with assays validated the accuracy of the models developed by my colleague Étienne."

The device allows for the simultaneous use of several reagents to detect various molecules in the same sample, which saves biologists valuable time. For certain types of tests, the analysis time could be reduced from several days to a few hours, or even a matter of minutes. In addition, the versatility of this technology should make it usable for various analytical processes, including immunological and DNA tests.


Professor Gervais's team is already considering a next step in his project: the development of a screen displaying a chemical image.

"It would be a sort of chemical equivalent of the liquid-crystal display," Professor Gervais explains. "In the same way that we move electrons across a screen, we would send jets of fluid at various concentrations that would react with a surface. Together, they would form an image. We are very excited to move forward with this project, for which we have obtained a provisional patent."


For now, the technology developed by this research team is aimed at the fundamental research market. "Our processes make it possible to expose cells to many reagents simultaneously," Professor Gervais says. "They can help biologists study the interactions between proteins and reagents on a large scale, increasing the amount and quality of information obtained during assays."

He explains that subsequently, the pharmaceutical market will also be able to benefit from new methods of screening-system automation resulting from the discovery. Lastly, it opens up a new avenue for drug discovery by facilitating patient cell culture and exposure to various drug agents to determine which ones they respond to best.


A biosensor typically consists of a bio-receptor (enzyme/antibody/cell/nucleic acid/aptamer), transducer component (semi-conducting material/nanomaterial), and electronic system which includes a signal amplifier, processor & display. [5] Transducers and electronics can be combined, e.g., in CMOS-based microsensor systems. [6] [7] The recognition component, often called a bioreceptor, uses biomolecules from organisms or receptors modeled after biological systems to interact with the analyte of interest. This interaction is measured by the biotransducer which outputs a measurable signal proportional to the presence of the target analyte in the sample. The general aim of the design of a biosensor is to enable quick, convenient testing at the point of concern or care where the sample was procured. [8] [9]

In a biosensor, the bioreceptor is designed to interact with the specific analyte of interest to produce an effect measurable by the transducer. High selectivity for the analyte among a matrix of other chemical or biological components is a key requirement of the bioreceptor. While the type of biomolecule used can vary widely, biosensors can be classified according to common types of bioreceptor interactions involving: antibody/antigen, [10] enzymes/ligands, nucleic acids/DNA, cellular structures/cells, or biomimetic materials. [11] [12]

Antibody/antigen interactions Edit

An immunosensor utilizes the very specific binding affinity of antibodies for a specific compound or antigen. The specific nature of the antibody-antigen interaction is analogous to a lock and key fit in that the antigen will only bind to the antibody if it has the correct conformation. Binding events result in a physicochemical change that in combination with a tracer, such as fluorescent molecules, enzymes, or radioisotopes, can generate a signal. There are limitations with using antibodies in sensors: 1. The antibody binding capacity is strongly dependent on assay conditions (e.g. pH and temperature), and 2. the antibody-antigen interaction is generally robust, however, binding can be disrupted by chaotropic reagents, organic solvents, or even ultrasonic radiation. [13]

Antibody-antigen interactions can also be used for serological testing, or the detection of circulating antibodies in response to a specific disease. Importantly, serology tests have become an important part of the global response to the Covid-19 pandemic. [14]

Artificial binding proteins Edit

The use of antibodies as the bio-recognition component of biosensors has several drawbacks. They have high molecular weights and limited stability, contain essential disulfide bonds and are expensive to produce. In one approach to overcome these limitations, recombinant binding fragments (Fab, Fv or scFv) or domains (VH, VHH) of antibodies have been engineered. [15] In another approach, small protein scaffolds with favorable biophysical properties have been engineered to generate artificial families of Antigen Binding Proteins (AgBP), capable of specific binding to different target proteins while retaining the favorable properties of the parent molecule. The elements of the family that specifically bind to a given target antigen, are often selected in vitro by display techniques: phage display, ribosome display, yeast display or mRNA display. The artificial binding proteins are much smaller than antibodies (usually less than 100 amino-acid residues), have a strong stability, lack disulfide bonds and can be expressed in high yield in reducing cellular environments like the bacterial cytoplasm, contrary to antibodies and their derivatives. [16] [17] They are thus especially suitable to create biosensors. [18] [19]

Enzymatic interactions Edit

The specific binding capabilities and catalytic activity of enzymes make them popular bioreceptors. Analyte recognition is enabled through several possible mechanisms: 1) the enzyme converting the analyte into a product that is sensor-detectable, 2) detecting enzyme inhibition or activation by the analyte, or 3) monitoring modification of enzyme properties resulting from interaction with the analyte. [13] The main reasons for the common use of enzymes in biosensors are: 1) ability to catalyze a large number of reactions 2) potential to detect a group of analytes (substrates, products, inhibitors, and modulators of the catalytic activity) and 3) suitability with several different transduction methods for detecting the analyte. Notably, since enzymes are not consumed in reactions, the biosensor can easily be used continuously. The catalytic activity of enzymes also allows lower limits of detection compared to common binding techniques. However, the sensor's lifetime is limited by the stability of the enzyme.

Affinity binding receptors Edit

Antibodies have a high binding constant in excess of 10^8 L/mol, which stands for a nearly irreversible association once the antigen-antibody couple has formed. For certain analyte molecules like glucose affinity binding proteins exist that bind their ligand with a high specificity like an antibody, but with a much smaller binding constant on the order of 10^2 to 10^4 L/mol. The association between analyte and receptor then is of reversible nature and next to the couple between both also their free molecules occur in a measurable concentration. In case of glucose, for instance, concanavalin A may function as affinity receptor exhibiting a binding constant of 4x10^2 L/mol. [20] The use of affinity binding receptors for purposes of biosensing has been proposed by Schultz and Sims in 1979 [21] and was subsequently configured into a fluorescent assay for measuring glucose in the relevant physiological range between 4.4 and 6.1 mmol/L. [22] The sensor principle has the advantage that it does not consume the analyte in a chemical reaction as occurs in enzymatic assays.

Nucleic acid interactions Edit

Biosensors employing nucleic acid based receptors can be either based on complementary base pairing interactions referred to as genosensors or specific nucleic acid based antibody mimics (aptamers) as aptasensors. [23] In the former, the recognition process is based on the principle of complementary base pairing, adenine:thymine and cytosine:guanine in DNA. If the target nucleic acid sequence is known, complementary sequences can be synthesized, labeled, and then immobilized on the sensor. The hybridization event can be optically detected and presence of target DNA/RNA ascertained. In the latter, aptamers generated against the target recognise it via interplay of specific non-covalent interactions and induced fitting. These aptamers can be labelled with a fluorophore/metal nanoparticles easily for optical detection or may be employed for label-free electrochemical or cantilever based detection platforms for a wide range of target molecules or complex targets like cells and viruses. [24] [25]

Epigenetics Edit

It has been proposed that properly optimized integrated optical resonators can be exploited for detecting epigenetic modifications (e.g. DNA methylation, histone post-translational modifications) in body fluids from patients affected by cancer or other diseases. [26] Photonic biosensors with ultra-sensitivity are nowadays being developed at a research level to easily detect cancerous cells within the patient's urine. [27] Different research projects aim to develop new portable devices that use cheap, environmentally friendly, disposable cartridges that require only simple handling with no need of further processing, washing, or manipulation by expert technicians. [28]

Organelles Edit

Organelles form separate compartments inside cells and usually perform function independently. Different kinds of organelles have various metabolic pathways and contain enzymes to fulfill its function. Commonly used organelles include lysosome, chloroplast and mitochondria. The spatial-temporal distribution pattern of calcium is closely related to ubiquitous signaling pathway. Mitochondria actively participate in the metabolism of calcium ions to control the function and also modulate the calcium related signaling pathways. Experiments have proved that mitochondria have the ability to respond to high calcium concentrations generated in their proximity by opening the calcium channels. [29] In this way, mitochondria can be used to detect the calcium concentration in medium and the detection is very sensitive due to high spatial resolution. Another application of mitochondria is used for detection of water pollution. Detergent compounds' toxicity will damage the cell and subcellular structure including mitochondria. The detergents will cause a swelling effect which could be measured by an absorbance change. Experiment data shows the change rate is proportional to the detergent concentration, providing a high standard for detection accuracy. [30]

Cells Edit

Cells are often used in bioreceptors because they are sensitive to surrounding environment and they can respond to all kinds of stimulants. Cells tend to attach to the surface so they can be easily immobilized. Compared to organelles they remain active for longer period and the reproducibility makes them reusable. They are commonly used to detect global parameter like stress condition, toxicity and organic derivatives. They can also be used to monitor the treatment effect of drugs. One application is to use cells to determine herbicides which are main aquatic contaminant. [31] Microalgae are entrapped on a quartz microfiber and the chlorophyll fluorescence modified by herbicides is collected at the tip of an optical fiber bundle and transmitted to a fluorimeter. The algae are continuously cultured to get optimized measurement. Results show that detection limit of certain herbicide can reach sub-ppb concentration level. Some cells can also be used to monitor the microbial corrosion. [32] Pseudomonas sp. is isolated from corroded material surface and immobilized on acetylcellulose membrane. The respiration activity is determined by measuring oxygen consumption. There is linear relationship between the current generated and the concentration of sulfuric acid. The response time is related to the loading of cells and surrounding environments and can be controlled to no more than 5min.

Tissue Edit

Tissues are used for biosensor for the abundance of enzymes existing. Advantages of tissues as biosensors include the following: [33]

  • easier to immobilize compared to cells and organelles
  • the higher activity and stability from maintaining enzymes in the natural environment
  • the availability and low price
  • the avoidance of tedious work of extraction, centrifuge, and purification of enzymes
  • necessary cofactors for an enzyme to function exists
  • the diversity providing a wide range of choices concerning different objectives.

There also exist some disadvantages of tissues, like the lack of specificity due to the interference of other enzymes and longer response time due to the transport barrier.

An important part of a biosensor is to attach the biological elements (small molecules/protein/cells) to the surface of the sensor (be it metal, polymer, or glass). The simplest way is to functionalize the surface in order to coat it with the biological elements. This can be done by polylysine, aminosilane, epoxysilane, or nitrocellulose in the case of silicon chips/silica glass. Subsequently, the bound biological agent may also be fixed—for example, by layer by layer deposition of alternatively charged polymer coatings. [34]

Alternatively, three-dimensional lattices (hydrogel/xerogel) can be used to chemically or physically entrap these (whereby chemically entrapped it is meant that the biological element is kept in place by a strong bond, while physically they are kept in place being unable to pass through the pores of the gel matrix). The most commonly used hydrogel is sol-gel, glassy silica generated by polymerization of silicate monomers (added as tetra alkyl orthosilicates, such as TMOS or TEOS) in the presence of the biological elements (along with other stabilizing polymers, such as PEG) in the case of physical entrapment. [35]

Another group of hydrogels, which set under conditions suitable for cells or protein, are acrylate hydrogel, which polymerizes upon radical initiation. One type of radical initiator is a peroxide radical, typically generated by combining a persulfate with TEMED (Polyacrylamide gel are also commonly used for protein electrophoresis), [36] alternatively light can be used in combination with a photoinitiator, such as DMPA (2,2-dimethoxy-2-phenylacetophenone). [37] Smart materials that mimic the biological components of a sensor can also be classified as biosensors using only the active or catalytic site or analogous configurations of a biomolecule. [38]

Biosensors can be classified by their biotransducer type. The most common types of biotransducers used in biosensors are:

  • electrochemical biosensors
  • optical biosensors
  • electronic biosensors
  • piezoelectric biosensors
  • gravimetric biosensors
  • pyroelectric biosensors
  • magnetic biosensors

Electrochemical Edit

Electrochemical biosensors are normally based on enzymatic catalysis of a reaction that produces or consumes electrons (such enzymes are rightly called redox enzymes). The sensor substrate usually contains three electrodes a reference electrode, a working electrode and a counter electrode. The target analyte is involved in the reaction that takes place on the active electrode surface, and the reaction may cause either electron transfer across the double layer (producing a current) or can contribute to the double layer potential (producing a voltage). We can either measure the current (rate of flow of electrons is now proportional to the analyte concentration) at a fixed potential or the potential can be measured at zero current (this gives a logarithmic response). Note that potential of the working or active electrode is space charge sensitive and this is often used. Further, the label-free and direct electrical detection of small peptides and proteins is possible by their intrinsic charges using biofunctionalized ion-sensitive field-effect transistors. [39]

Another example, the potentiometric biosensor, (potential produced at zero current) gives a logarithmic response with a high dynamic range. Such biosensors are often made by screen printing the electrode patterns on a plastic substrate, coated with a conducting polymer and then some protein (enzyme or antibody) is attached. They have only two electrodes and are extremely sensitive and robust. They enable the detection of analytes at levels previously only achievable by HPLC and LC/MS and without rigorous sample preparation. All biosensors usually involve minimal sample preparation as the biological sensing component is highly selective for the analyte concerned. The signal is produced by electrochemical and physical changes in the conducting polymer layer due to changes occurring at the surface of the sensor. Such changes can be attributed to ionic strength, pH, hydration and redox reactions, the latter due to the enzyme label turning over a substrate. [40] Field effect transistors, in which the gate region has been modified with an enzyme or antibody, can also detect very low concentrations of various analytes as the binding of the analyte to the gate region of the FET cause a change in the drain-source current.

Impedance spectroscopy based biosensor development has been gaining traction nowadays and many such devices / developments are found in the academia and industry. One such device, based on a 4-electrode electrochemical cell, using a nanoporous alumina membrane, has been shown to detect low concentrations of human alpha thrombin in presence of high background of serum albumin. [41] Also interdigitated electrodes have been used for impedance biosensors. [42]

Ion channel switch Edit

The use of ion channels has been shown to offer highly sensitive detection of target biological molecules. [43] By embedding the ion channels in supported or tethered bilayer membranes (t-BLM) attached to a gold electrode, an electrical circuit is created. Capture molecules such as antibodies can be bound to the ion channel so that the binding of the target molecule controls the ion flow through the channel. This results in a measurable change in the electrical conduction which is proportional to the concentration of the target.

An ion channel switch (ICS) biosensor can be created using gramicidin, a dimeric peptide channel, in a tethered bilayer membrane. [44] One peptide of gramicidin, with attached antibody, is mobile and one is fixed. Breaking the dimer stops the ionic current through the membrane. The magnitude of the change in electrical signal is greatly increased by separating the membrane from the metal surface using a hydrophilic spacer.

Quantitative detection of an extensive class of target species, including proteins, bacteria, drug and toxins has been demonstrated using different membrane and capture configurations. [45] [46] The European research project Greensense develops a biosensor to perform quantitative screening of drug-of-abuse such as THC, morphine, and cocaine [47] in saliva and urine.

Reagentless fluorescent biosensor Edit

A reagentless biosensor can monitor a target analyte in a complex biological mixture without additional reagent. Therefore, it can function continuously if immobilized on a solid support. A fluorescent biosensor reacts to the interaction with its target analyte by a change of its fluorescence properties. A Reagentless Fluorescent biosensor (RF biosensor) can be obtained by integrating a biological receptor, which is directed against the target analyte, and a solvatochromic fluorophore, whose emission properties are sensitive to the nature of its local environment, in a single macromolecule. The fluorophore transduces the recognition event into a measurable optical signal. The use of extrinsic fluorophores, whose emission properties differ widely from those of the intrinsic fluorophores of proteins, tryptophan and tyrosine, enables one to immediately detect and quantify the analyte in complex biological mixtures. The integration of the fluorophore must be done in a site where it is sensitive to the binding of the analyte without perturbing the affinity of the receptor.

Antibodies and artificial families of Antigen Binding Proteins (AgBP) are well suited to provide the recognition module of RF biosensors since they can be directed against any antigen (see the paragraph on bioreceptors). A general approach to integrate a solvatochromic fluorophore in an AgBP when the atomic structure of the complex with its antigen is known, and thus transform it into a RF biosensor, has been described. [18] A residue of the AgBP is identified in the neighborhood of the antigen in their complex. This residue is changed into a cysteine by site-directed mutagenesis. The fluorophore is chemically coupled to the mutant cysteine. When the design is successful, the coupled fluorophore does not prevent the binding of the antigen, this binding shields the fluorophore from the solvent, and it can be detected by a change of fluorescence. This strategy is also valid for antibody fragments. [48] [49]

However, in the absence of specific structural data, other strategies must be applied. Antibodies and artificial families of AgBPs are constituted by a set of hypervariable (or randomized) residue positions, located in a unique sub-region of the protein, and supported by a constant polypeptide scaffold. The residues that form the binding site for a given antigen, are selected among the hypervariable residues. It is possible to transform any AgBP of these families into a RF biosensor, specific of the target antigen, simply by coupling a solvatochromic fluorophore to one of the hypervariable residues that have little or no importance for the interaction with the antigen, after changing this residue into cysteine by mutagenesis. More specifically, the strategy consists in individually changing the residues of the hypervariable positions into cysteine at the genetic level, in chemically coupling a solvatochromic fluorophore with the mutant cysteine, and then in keeping the resulting conjugates that have the highest sensitivity (a parameter that involves both affinity and variation of fluorescence signal). [19] This approach is also valid for families of antibody fragments. [50]

A posteriori studies have shown that the best reagentless fluorescent biosensors are obtained when the fluorophore does not make non-covalent interactions with the surface of the bioreceptor, which would increase the background signal, and when it interacts with a binding pocket at the surface of the target antigen. [51] The RF biosensors that are obtained by the above methods, can function and detect target analytes inside living cells. [52]

Magnetic biosensors Edit

Magnetic biosensors utilize paramagnetic or supra-paramagnetic particles, or crystals, to detect biological interactions. Examples could be coil-inductance, resistance, or other magnetic properties. It is common to use magnetic nano or microparticles. In the surface of such particles are the bioreceptors, that can be DNA (complementary to a sequence or aptamers) antibodies, or others. The binding of the bioreceptor will affect some of the magnetic particle properties that can be measured by AC susceptometry, [53] a Hall Effect sensor, [54] a giant magnetoresistance device, [55] or others.

Others Edit

Piezoelectric sensors utilise crystals which undergo an elastic deformation when an electrical potential is applied to them. An alternating potential (A.C.) produces a standing wave in the crystal at a characteristic frequency. This frequency is highly dependent on the elastic properties of the crystal, such that if a crystal is coated with a biological recognition element the binding of a (large) target analyte to a receptor will produce a change in the resonance frequency, which gives a binding signal. In a mode that uses surface acoustic waves (SAW), the sensitivity is greatly increased. This is a specialised application of the quartz crystal microbalance as a biosensor

Electrochemiluminescence (ECL) is nowadays a leading technique in biosensors. [56] [57] [58] Since the excited species are produced with an electrochemical stimulus rather than with a light excitation source, ECL displays improved signal-to-noise ratio compared to photoluminescence, with minimized effects due to light scattering and luminescence background. In particular, coreactant ECL operating in buffered aqueous solution in the region of positive potentials (oxidative-reduction mechanism) definitively boosted ECL for immunoassay, as confirmed by many research applications and, even more, by the presence of important companies which developed commercial hardware for high throughput immunoassays analysis in a market worth billions of dollars each year.

Thermometric biosensors are rare.

The MOSFET (metal-oxide-semiconductor field-effect transistor, or MOS transistor) was invented by Mohamed M. Atalla and Dawon Kahng in 1959, and demonstrated in 1960. [59] Two years later, Leland C. Clark and Champ Lyons invented the first biosensor in 1962. [60] [61] Biosensor MOSFETs (BioFETs) were later developed, and they have since been widely used to measure physical, chemical, biological and environmental parameters. [62]

The first BioFET was the ion-sensitive field-effect transistor (ISFET), invented by Piet Bergveld for electrochemical and biological applications in 1970. [63] [64] the adsorption FET (ADFET) was patented by P.F. Cox in 1974, and a hydrogen-sensitive MOSFET was demonstrated by I. Lundstrom, M.S. Shivaraman, C.S. Svenson and L. Lundkvist in 1975. [62] The ISFET is a special type of MOSFET with a gate at a certain distance, [62] and where the metal gate is replaced by an ion-sensitive membrane, electrolyte solution and reference electrode. [65] The ISFET is widely used in biomedical applications, such as the detection of DNA hybridization, biomarker detection from blood, antibody detection, glucose measurement, pH sensing, and genetic technology. [65]

By the mid-1980s, other BioFETs had been developed, including the gas sensor FET (GASFET), pressure sensor FET (PRESSFET), chemical field-effect transistor (ChemFET), reference ISFET (REFET), enzyme-modified FET (ENFET) and immunologically modified FET (IMFET). [62] By the early 2000s, BioFETs such as the DNA field-effect transistor (DNAFET), gene-modified FET (GenFET) and cell-potential BioFET (CPFET) had been developed. [65]

The appropriate placement of biosensors depends on their field of application, which may roughly be divided into biotechnology, agriculture, food technology and biomedicine.

In biotechnology, analysis of the chemical composition of cultivation broth can be conducted in-line, on-line, at-line and off-line. As outlined by the US Food and Drug Administration (FDA) the sample is not removed from the process stream for in-line sensors, while it is diverted from the manufacturing process for on-line measurements. For at-line sensors the sample may be removed and analyzed in close proximity to the process stream. [66] An example of the latter is the monitoring of lactose in a dairy processing plant. [67] Off-line biosensors compare to bioanalytical techniques that are not operating in the field, but in the laboratory. These techniques are mainly used in agriculture, food technology and biomedicine.

In medical applications biosensors are generally categorized as in vitro and in vivo systems. An in vitro, biosensor measurement takes place in a test tube, a culture dish, a microtiter plate or elsewhere outside a living organism. The sensor uses a bioreceptor and transducer as outlined above. An example of an in vitro biosensor is an enzyme-conductimetric biosensor for blood glucose monitoring. There is a challenge to create a biosensor that operates by the principle of point-of-care testing, i.e. at the location where the test is needed. [68] [69] Development of wearable biosensors is among such studies. [70] The elimination of lab testing can save time and money. An application of a POCT biosensor can be for the testing of HIV in areas where it is difficult for patients to be tested. A biosensor can be sent directly to the location and a quick and easy test can be used.

An in vivo biosensor is an implantable device that operates inside the body. Of course, biosensor implants have to fulfill the strict regulations on sterilization in order to avoid an initial inflammatory response after implantation. The second concern relates to the long-term biocompatibility, i.e. the unharmful interaction with the body environment during the intended period of use. [72] Another issue that arises is failure. If there is failure, the device must be removed and replaced, causing additional surgery. An example for application of an in vivo biosensor would be the insulin monitoring within the body, which is not available yet.

Most advanced biosensor implants have been developed for the continuous monitoring of glucose. [73] [74] The figure displays a device, for which a Ti casing and a battery as established for cardiovascular implants like pacemakers and defibrillators is used. [71] Its size is determined by the battery as required for a lifetime of one year. Measured glucose data will be transmitted wirelessly out of the body within the MICS 402-405 MHz band as approved for medical implants.

Biosensors can also be integrated into mobile phone systems, making them user-friendly and accessible to a large number of users. [75]

There are many potential applications of biosensors of various types. The main requirements for a biosensor approach to be valuable in terms of research and commercial applications are the identification of a target molecule, availability of a suitable biological recognition element, and the potential for disposable portable detection systems to be preferred to sensitive laboratory-based techniques in some situations. Some examples are glucose monitoring in diabetes patients, other medical health related targets, environmental applications, e.g. the detection of pesticides and river water contaminants, such as heavy metal ions, [76] remote sensing of airborne bacteria, e.g. in counter-bioterrorist activities, remote sensing of water quality in coastal waters by describing online different aspects of clam ethology (biological rhythms, growth rates, spawning or death records) in groups of abandoned bivalves around the world, [77] detection of pathogens, determining levels of toxic substances before and after bioremediation, detection and determining of organophosphate, routine analytical measurement of folic acid, biotin, vitamin B12 and pantothenic acid as an alternative to microbiological assay, determination of drug residues in food, such as antibiotics and growth promoters, particularly meat and honey, drug discovery and evaluation of biological activity of new compounds, protein engineering in biosensors, [78] and detection of toxic metabolites such as mycotoxins.

A common example of a commercial biosensor is the blood glucose biosensor, which uses the enzyme glucose oxidase to break blood glucose down. In doing so it first oxidizes glucose and uses two electrons to reduce the FAD (a component of the enzyme) to FADH2. This in turn is oxidized by the electrode in a number of steps. The resulting current is a measure of the concentration of glucose. In this case, the electrode is the transducer and the enzyme is the biologically active component.

A canary in a cage, as used by miners to warn of gas, could be considered a biosensor. Many of today's biosensor applications are similar, in that they use organisms which respond to toxic substances at a much lower concentrations than humans can detect to warn of their presence. Such devices can be used in environmental monitoring, [77] trace gas detection and in water treatment facilities.

Many optical biosensors are based on the phenomenon of surface plasmon resonance (SPR) techniques. [79] [80] This utilises a property of and other materials specifically that a thin layer of gold on a high refractive index glass surface can absorb laser light, producing electron waves (surface plasmons) on the gold surface. This occurs only at a specific angle and wavelength of incident light and is highly dependent on the surface of the gold, such that binding of a target analyte to a receptor on the gold surface produces a measurable signal.

Surface plasmon resonance sensors operate using a sensor chip consisting of a plastic cassette supporting a glass plate, one side of which is coated with a microscopic layer of gold. This side contacts the optical detection apparatus of the instrument. The opposite side is then contacted with a microfluidic flow system. The contact with the flow system creates channels across which reagents can be passed in solution. This side of the glass sensor chip can be modified in a number of ways, to allow easy attachment of molecules of interest. Normally it is coated in carboxymethyl dextran or similar compound.

The refractive index at the flow side of the chip surface has a direct influence on the behavior of the light reflected off the gold side. Binding to the flow side of the chip has an effect on the refractive index and in this way biological interactions can be measured to a high degree of sensitivity with some sort of energy. The refractive index of the medium near the surface changes when biomolecules attach to the surface, and the SPR angle varies as a function of this change.

Light of a fixed wavelength is reflected off the gold side of the chip at the angle of total internal reflection, and detected inside the instrument. The angle of incident light is varied in order to match the evanescent wave propagation rate with the propagation rate of the surface plasmon plaritons. [81] This induces the evanescent wave to penetrate through the glass plate and some distance into the liquid flowing over the surface.

Other optical biosensors are mainly based on changes in absorbance or fluorescence of an appropriate indicator compound and do not need a total internal reflection geometry. For example, a fully operational prototype device detecting casein in milk has been fabricated. The device is based on detecting changes in absorption of a gold layer. [82] A widely used research tool, the micro-array, can also be considered a biosensor.

Biological biosensors often incorporate a genetically modified form of a native protein or enzyme. The protein is configured to detect a specific analyte and the ensuing signal is read by a detection instrument such as a fluorometer or luminometer. An example of a recently developed biosensor is one for detecting cytosolic concentration of the analyte cAMP (cyclic adenosine monophosphate), a second messenger involved in cellular signaling triggered by ligands interacting with receptors on the cell membrane. [83] Similar systems have been created to study cellular responses to native ligands or xenobiotics (toxins or small molecule inhibitors). Such "assays" are commonly used in drug discovery development by pharmaceutical and biotechnology companies. Most cAMP assays in current use require lysis of the cells prior to measurement of cAMP. A live-cell biosensor for cAMP can be used in non-lysed cells with the additional advantage of multiple reads to study the kinetics of receptor response.

Nanobiosensors use an immobilized bioreceptor probe that is selective for target analyte molecules. Nanomaterials are exquisitely sensitive chemical and biological sensors. Nanoscale materials demonstrate unique properties. Their large surface area to volume ratio can achieve rapid and low cost reactions, using a variety of designs. [84]

Other evanescent wave biosensors have been commercialised using waveguides where the propagation constant through the waveguide is changed by the absorption of molecules to the waveguide surface. One such example, dual polarisation interferometry uses a buried waveguide as a reference against which the change in propagation constant is measured. Other configurations such as the Mach–Zehnder have reference arms lithographically defined on a substrate. Higher levels of integration can be achieved using resonator geometries where the resonant frequency of a ring resonator changes when molecules are absorbed. [85] [86]

Recently, arrays of many different detector molecules have been applied in so called electronic nose devices, where the pattern of response from the detectors is used to fingerprint a substance. [87] In the Wasp Hound odor-detector, the mechanical element is a video camera and the biological element is five parasitic wasps who have been conditioned to swarm in response to the presence of a specific chemical. [88] Current commercial electronic noses, however, do not use biological elements.

Glucose monitoring Edit

Commercially available glucose monitors rely on amperometric sensing of glucose by means of glucose oxidase, which oxidises glucose producing hydrogen peroxide which is detected by the electrode. To overcome the limitation of amperometric sensors, a flurry of research is present into novel sensing methods, such as fluorescent glucose biosensors. [89]

Interferometric reflectance imaging sensor Edit

The interferometric reflectance imaging sensor (IRIS) is based on the principles of optical interference and consists of a silicon-silicon oxide substrate, standard optics, and low-powered coherent LEDs. When light is illuminated through a low magnification objective onto the layered silicon-silicon oxide substrate, an interferometric signature is produced. As biomass, which has a similar index of refraction as silicon oxide, accumulates on the substrate surface, a change in the interferometric signature occurs and the change can be correlated to a quantifiable mass. Daaboul et al. used IRIS to yield a label-free sensitivity of approximately 19 ng/mL. [90] Ahn et al. improved the sensitivity of IRIS through a mass tagging technique. [91]

Since initial publication, IRIS has been adapted to perform various functions. First, IRIS integrated a fluorescence imaging capability into the interferometric imaging instrument as a potential way to address fluorescence protein microarray variability. [92] Briefly, the variation in fluorescence microarrays mainly derives from inconsistent protein immobilization on surfaces and may cause misdiagnoses in allergy microarrays. [93] To correct for any variation in protein immobilization, data acquired in the fluorescence modality is then normalized by the data acquired in the label-free modality. [93] IRIS has also been adapted to perform single nanoparticle counting by simply switching the low magnification objective used for label-free biomass quantification to a higher objective magnification. [94] [95] This modality enables size discrimination in complex human biological samples. Monroe et al. used IRIS to quantify protein levels spiked into human whole blood and serum and determined allergen sensitization in characterized human blood samples using zero sample processing. [96] Other practical uses of this device include virus and pathogen detection. [97]

Food analysis Edit

There are several applications of biosensors in food analysis. In the food industry, optics coated with antibodies are commonly used to detect pathogens and food toxins. Commonly, the light system in these biosensors is fluorescence, since this type of optical measurement can greatly amplify the signal.

A range of immuno- and ligand-binding assays for the detection and measurement of small molecules such as water-soluble vitamins and chemical contaminants (drug residues) such as sulfonamides and Beta-agonists have been developed for use on SPR based sensor systems, often adapted from existing ELISA or other immunological assay. These are in widespread use across the food industry.

DNA biosensors Edit

DNA can be the analyte of a biosensor, being detected through specific means, but it can also be used as part of a biosensor or, theoretically, even as a whole biosensor.

Many techniques exist to detect DNA, which is usually a means to detect organisms that have that particular DNA. DNA sequences can also be used as described above. But more forward-looking approaches exist, where DNA can be synthesized to hold enzymes in a biological, stable gel. [98] Other applications are the design of aptamers, sequences of DNA that have a specific shape to bind a desired molecule. The most innovative processes use DNA origami for this, creating sequences that fold in a predictable structure that is useful for detection. [99] [100]

Microbial biosensors Edit

Microbial biosensors exploit the response of bacteria to a given substance. For example, arsenic can be detected using the ars operon found in several bacterial taxon. [101]

Ozone biosensors Edit

Because ozone filters out harmful ultraviolet radiation, the discovery of holes in the ozone layer of the earth's atmosphere has raised concern about how much ultraviolet light reaches the earth's surface. Of particular concern are the questions of how deeply into sea water ultraviolet radiation penetrates and how it affects marine organisms, especially plankton (floating microorganisms) and viruses that attack plankton. Plankton form the base of the marine food chains and are believed to affect our planet's temperature and weather by uptake of CO2 for photosynthesis.

Deneb Karentz, a researcher at the Laboratory of Radio-biology and Environmental Health (University of California, San Francisco) has devised a simple method for measuring ultraviolet penetration and intensity. Working in the Antarctic Ocean, she submerged to various depths thin plastic bags containing special strains of E. coli that are almost totally unable to repair ultraviolet radiation damage to their DNA. Bacterial death rates in these bags were compared with rates in unexposed control bags of the same organism. The bacterial "biosensors" revealed constant significant ultraviolet damage at depths of 10 m and frequently at 20 and 30 m. Karentz plans additional studies of how ultraviolet may affect seasonal plankton blooms (growth spurts) in the oceans. [102]

Metastatic cancer cell biosensors Edit

Metastasis is the spread of cancer from one part of the body to another via either the circulatory system or lymphatic system. [103] Unlike radiology imaging tests (mammograms), which send forms of energy (x-rays, magnetic fields, etc.) through the body to only take interior pictures, biosensors have the potential to directly test the malignant power of the tumor. The combination of a biological and detector element allows for a small sample requirement, a compact design, rapid signals, rapid detection, high selectivity and high sensitivity for the analyte being studied. Compared to the usual radiology imaging tests biosensors have the advantage of not only finding out how far cancer has spread and checking if treatment is effective but also are cheaper, more efficient (in time, cost and productivity) ways to assess metastaticity in early stages of cancer.

Biological engineering researchers have created oncological biosensors for breast cancer. [104] Breast cancer is the leading common cancer among women worldwide. [105] An example would be a transferrin- quartz crystal microbalance (QCM). As a biosensor, quartz crystal microbalances produce oscillations in the frequency of the crystal's standing wave from an alternating potential to detect nano-gram mass changes. These biosensors are specifically designed to interact and have high selectivity for receptors on cell (cancerous and normal) surfaces. Ideally, this provides a quantitative detection of cells with this receptor per surface area instead of a qualitative picture detection given by mammograms.

Seda Atay, a biotechnology researcher at Hacettepe University, experimentally observed this specificity and selectivity between a QCM and MDA-MB 231 breast cells, MCF 7 cells, and starved MDA-MB 231 cells in vitro. [104] With other researchers she devised a method of washing these different metastatic leveled cells over the sensors to measure mass shifts due to different quantities of transferrin receptors. Particularly, the metastatic power of breast cancer cells can be determined by Quartz crystal microbalances with nanoparticles and transferrin that would potentially attach to transferrin receptors on cancer cell surfaces. There is very high selectivity for transferrin receptors because they are over-expressed in cancer cells. If cells have high expression of transferrin receptors, which shows their high metastatic power, they have higher affinity and bind more to the QCM that measures the increase in mass. Depending on the magnitude of the nano-gram mass change, the metastatic power can be determined.

Additionally, in the last years, significant attentions have been focused to detect the biomarkers of lung cancer without biopsy. In this regard, biosensors are very attractive and applicable tools for providing rapid, sensitive, specific, stable, cost-effective and non-invasive detections for early lung cancer diagnosis. Thus, cancer biosensors consisting of specific biorecognition molecules such as antibodies, complementary nucleic acid probes or other immobilized biomolecules on a transducer surface. The biorecognition molecules interact specifically with the biomarkers (targets) and the generated biological responses are converted by the transducer into a measurable analytical signal. Depending on the type of biological response, various transducers are utilized in the fabrication of cancer biosensors such as electrochemical, optical and mass-based transducers. [106]

Experimental Results

When the chip is oriented on its edge relative to gravity in Fig 3A (θ = 0°), the more-dense blue fluid remains on the bottom of the horizontal channel and the less-dense yellow fluid remains on the top of the channel. Consequently, the fluids exit the mixer chip in the same directions from which they entered (yellow fluid at Outlet 1 and blue fluid at Outlet 2). This case is identical to the simulation shown in Fig 1A. When the chip is held flat relative to gravity in Fig 3B (θ = 90°), the two fluids reorient relative to gravity (rotating 90° clockwise) to place the more-dense blue fluid at the bottom of the horizontal channel and the less-dense yellow fluid on the top of the channel. When the horizontal channel splits, both output channels have the same contents (∼50% yellow and ∼50% blue, which appears as green in the outlets). This situation is identical to the simulation shown in Fig 2C. Finally, when the chip is oriented on its other edge in Fig 3C (θ = 180°), the yellow and blue fluids swap places in the horizontal channel to place the more-dense blue fluid on the bottom and the less-dense yellow fluid on the top. Consequently, the fluids leave the Outlet channels opposite from where they entered, with yellow fluid at Outlet 2 and blue fluid at Outlet 1.

To demonstrate orientation-based control of fluid flow over a wide range of angles (not just the three angles shown in Fig 3), the mixer chip was operated at angles from 0° to 180° in 15° increments. As before, the more-dense blue fluid was a sucrose solution (density = 1.07 g/mL) and the less-dense yellow fluid was water (density = 1.00 g/mL). Fig 4 shows the concentration of each dye in the two output channels. As the mixer chip’s angle of rotation θ is varied from 0° to 180°, the concentration of blue fluid rises and yellow fluid drops in Outlet 1, and an opposite trend is observed in Outlet 2. The relationship of concentration on angle of rotation is roughly linear as predicted by Eq 5. This experimental data deviates from the predicted model at angles near 0° and 180°, where the outlet fluids concentrations are not 100% and 0% as predicted but ∼90% and ∼10% instead. This is likely due to diffusion within the horizontal channel in the mixer chip, which contributes a small amount of mixing between the two fluid streams. The effect of this diffusional mixing is most pronounced at angles near 0° and 180° where the outlets should contain pure (unmixed) fluids according to Eq 5 instead their contents are ∼90% pure. Additionally, the relationship of fluid concentration on angle of rotation in Fig 4 is not purely linear but appears to have some higher-order shape. This can be attributed to the cross-sectional geometry of the 3D-printed microfluidic channel because of the limited resolution of the 3D printer. The 1 mm channel was printed using a stereolithography 3D printer with a tolerance of ± 200 μm. This could result in a microfluidic channel that is not perfectly circular and requires a more complex model than the one shown in Eq 5. However, the error bars in Fig 4 confirm that the outlet concentrations are a reproducible and predictable function of the angle of rotation of the chip, and a higher-order function could easily be derived that predicts the fluid concentrations in a chip at any rotation angle θ to within a few percentage points.

Also shown are photographs of the fluid collected at each outlet and an illustration of the chip’s orientation at each angle of rotation. N = 3 measurements for each point error bars indicate ±1 standard deviation. These results show that mixture composition is a function of the angle of rotation of the chip, and any desired mixture can be generated simply by orienting the chip and the required angle.

To explore the role of channel cross-section shape in orientation-controlled microfluidics, we designed and 3D printed mixer chips with square (1 × 1 mm) and rectangular (1 × 1.25 mm) cross-section channels. We then repeated the experiments from Fig 5A using the square and rectangular chips for rotation angles θ = 0°, 90°, and 180°. The results in Fig 5A confirm that the orientation of a chip can still be used to control fluid mixing in chips with square and rectangular cross sections, though increased deviation from the predicted fluid concentrations is observed in the rectangular cross-section chip. This suggests that orientation-based control of microfluidics works best in channels with aspect ratios near one at higher aspect ratios the channel geometry hinders the desired rotation of fluid in the channel. We were limited to a 1 mm channel diameter since we wanted to explore different geometries and aspect ratios using a 3D printer however, the same phenomena were observed in a conventional glass microfluidic channel of 180 μm (data not shown). We also examined the behavior of two fluids of different densities flowing in rectangular channels with much higher aspect ratios (cross-sectional dimensions 1 mm × 5 mm data not shown). The experimental results further support the assertion that orientation-based control of microfluidics is most practical in channels with cross-sectional aspect ratios near one (circular and square channels).

(A) Concentrations of yellow and blue dyes at Outlet 1 in mixer chips held at 0°, 90° and 180°, for chips with circular, square, and rectangular channel cross-sections. Chip orientation can be used to control on-chip fluids in devices with a variety of channel cross-sectional shapes, although performance deteriorates somewhat in the rectangular channel. These results suggest that microfluidic channels with aspect ratios close to one are best suited for orientation-based control. (B) Outlet fluid concentrations measured at Outlet 1 while orienting the mixer chip at 0°, 90° and 180° for four different fluid densities at Inlet 1. The density of fluid at Inlet 2 was kept constant at 1.00 g/mL. When there is no difference in fluid densities between the two inlet fluids, fluid reorientation is not observed and orientation-based control cannot be used (Control case). However, if the fluid density in Inlet 1 is just 2% greater than the fluid density in Inlet 2, then the flowing fluids reorient with respect to gravity and orientation-based control is possible.

In addition to channel geometry, fluid density has an effect on our phenomenon. The speed at which rotation occurs is influenced by the density difference. However, in practice we found that all rotations we observed were nearly instantaneous. The channel length was kept excessively long for characterization purposes. However, for a channel diameter of 1 mm, a minimum length of 10 mm is required for the rotation to occur. The larger the difference between fluid densities, the quicker the rotation occurs. Rotation speed could be increased by further increasing the difference between fluid densities however, the resulting increase in concentration gradients will result in an increase in diffusive flux, potentially allowing for more mixing between the two layers.

To determine how much of a difference in density between the two fluids is necessary to employ orientation-based control, sucrose solutions with different concentrations (and thus different densities) were prepared as shown in Fig 5B. Experiments were performed using solutions with densities of 1.00 g/mL (control), 1.02 g/mL, 1.07 g/mL, and 1.12 g/mL. As before, the density of the second fluid was kept constant at 1.00 g/mL (water). In the control case in Fig 5B, there is no difference in densities between the two fluids in the mixer chip, so no reorientation of the fluids occur within the chip and the fluid concentrations at the outlets are constant regardless of the angle of orientation of the chip however, when the density of one input fluid is increased by only 2% to 1.02 g/mL, the fluids again reorient with respect to gravity and orientation-based control is demonstrated.

Pneumatic Valves vs Peristaltic Pump

Alternatively, monolithic elastomeric valves and pumps can be fabricated using soft lithography for use on the microscale. An example of a elastomeric peristaltic pump can be seen to the right in Figure 3. Devices like this one, and other microfluidic devices, avoid several problems of alternative devices like the formation of bubbles and being dependent on flow. This design was arranged to have a crossed channel layout. The device works by pressure being applied to the upper channel which then causes the lower channel to close. This pushes the fluid along the pathway just like a peristaltic pump would. The fluid can remains sterile as it never needs to exit the tube or be exposed to outside contamination. The response time is extremely quick and the devices can be packed very close to each other. These pneumatic valves and pumps are useful for a large variety of applications, one of them being lab-on-a-chip [3].

Similar Records in DOE PAGES and OSTI.GOV collections:

3 million patients 2008. Despite the success of positron labeled tracers in many sciences, there is limited access in an affordable and convenient manner to develop and use new tracers. Integrated microfluidic chips are a new technology well matched to the concentrations of tracers. Our goal is to develop microfluidic chips and new synthesis approaches to enable wide dissemination of diverse types of tracers at low cost, and to produce new generations of radiochemists for which there are manymore » unfilled jobs. The program objectives are to: 1. Develop an integrated microfluidic platform technology for synthesizing and 18F-labeling diverse arrays of different classes of molecules. 2. Incorporate microfluidic chips into small PC controlled devices (“Synthesizer”) with a platform interfaced to PC for electronic and fluid input/out control. 3. Establish a de-centralized model with Synthesizers for discovering and producing molecular imaging probes, only requiring delivery of inexpensive [18F]fluoride ion from commercial PET radiopharmacies vs the centralized approach of cyclotron facilities synthesizing and shipping a few different types of 18F-probes. 4. Develop a position sensitive avalanche photo diode (PSAPD) camera for beta particles embedded in a microfluidic chip for imaging and measuring transport and biochemical reaction rates to valid new 18F-labeled probes in an array of cell cultures. These objectives are met within a research and educational program integrating radio-chemistry, synthetic chemistry, biochemistry, engineering and biology in the Crump Institute for Molecular Imaging. The Radiochemistry Training Program exposes PhD and post doctoral students to molecular imaging in vitro in cells and microorganisms in microfluidic chips and in vivo with PET, from new technologies for radiochemistry (macro to micro levels), biochemistry and biology to imaging principles, tracer kinetics, pharmacokinetics and biochemical assays. New generations of radiochemists will be immersed in the biochemistry and biology for which their labeled probes are being developed for assays of these processes. In this program engineers and radio-chemists integrate the principles of microfluidics and radiolabeling along with proper system design and chemistry rule sets to yield Synthesizers enabling biological and pharmaceutical scientists to develop diverse arrays of probes to pursue their interests. This progression would allow also radiochemists to focus on the further evolution of rapid, high yield synthetic reactions with new enabling technologies, rather than everyday production of radiotracers that should be done by technologists. The invention of integrated circuits in electronics established a platform technology that allowed an evolution of ideas and applications far beyond what could have been imagined at the beginning. Rather than provide a technology for the solution to a single problem, it is hoped that microfluidic radiochemistry will be an enabling platform technology for others to solve many problems. As part of this objective, another program goal is to commercialize the technologies that come from this work so that they can be provided to others who wish to use it. « less

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In a low-side switch, shown on the left, the load is between the power rail and the N-channel MOSFET doing the switching.

In a high-side switch, shown on the right, the load is between ground and the P-channel MOSFET doing the switching.

The low-side switches are convenient for driving LEDs, relays, motors etc. because you can generally driven them directly from a microcontroller's output, as long as the VWhat is the difference between using an open or a closed configuration for a microfluidic chip? - Biology,[nobr][H1toH2]

Watch the video: Live Demo of simple Microfluidic chip working. (November 2022).