There exists an incredible wealth of different techniques available to investigate ligand-receptor interaction from dialysis and isothermal titration calorimetry to electrophoresis and high-performance affinity chromatography. However, the focus will be placed here on surface-based biosensor technologies that allow for the direct measurement of ligand binding, since compared to other techniques, they offer the possibility of rapidly screening multiple recognition events. Furthermore, they may be combined with microfluidic han-dling, which makes them compatible with small sample volumes and thus ideally suited for studies of substances that are rare or time-consuming and expensive to obtain [7].
Current binding assays use surface-immobilised high-affinity capture ligands arranged in measurement chambers for parallel screening. The systems predominantly utilise either antibodies or proteins/peptides as binder molecules to capture circulating ligands during incubation [8]. Unbound ligands are removed by rinsing, and captured ligands can be detected via various investigation methods. Binding assays may be categorised according to their detection method.
1.1.1. Optical Sensors
Many methodologies currently used in optical biosensors require modification of one or more of the reaction components with labels (i. e. fluorophores). These approaches can have certain experimental limitations. For example, labels may directly or indirectly affect the binding of the reaction components [9–11]. Nonetheless, the popularity of these assay technologies is profound, since their sensitivity is still unmatched by label-free technologies.
Fluorescence Based Sensing Schemes
Fluorescence based biosensors usually function in a way that the specific formation of a noncovalent complex yields a fluorescence readout, which is meant to indicate the state or abundance of a particular target [9]. Irrespective of the molecular details, intracellular biosensors may be classified as either intramolecular, where the molecular recognition element and its target are contained within the same chain (connected by a flexible linker), or intermolecular, where the recognition element binds to form a bimolecular complex with a target that is endogenous to the cell [12]. Biosensors of the first type include those based on intramolecular Förster resonance energy transfer (FRET), with donor and
1.1 Biosensors for Molecular Recognition Events
acceptor fluorophores flanking the two ends of the chain [9,12]. Intermolecular biosensors include those based on membrane translocation or solvent-sensitive fluorescence [13,14].
Of these methods, FRET is the most prominent technique. FRET relies on the energy transfer between two chromophores for sensing purposes. The efficiency of the Förster resonance energy transfer between two fluorophores depends on their relative geometry (distance and orientation) as well as their spectral properties. The Förster distance is characteristic of the pair of fluorophores used as donor/acceptor and their relative orien-tation expressed by an orienorien-tation factor. This factor describes the angular dependence of the energy transfer. It is maximal when the dipoles of the fluorophores are collinear and becomes zero when they are orthogonal. As the Förster distance is usually in the order of 5 - 10 nm, nanometer scale conformational changes can be read out through changes in the FRET efficiency. The methods most widely applied to read-out FRET efficiency are intensity-based or fluorescence lifetime-based measurements [15].
1.1.2. Label-free Sensing Schemes
Label-free technologies offer a number of distinct advantages over label-dependent assay formats. First, they are non-invasive and require minimal manipulation of reaction com-ponents, such as proteins or cells, thus enhancing the potential for measuring biologically meaningful data [10]. And second, label-free methods do not suffer from potential assay artifacts such as compound autofluorescence or quenching as no fluorescent dye or label is involved [10]. Lable-free biosensing schemes may be based on methods such as ellip-sometry, surface plasmon resonance (SPR) spectroscopy, waveguides and reflectometric interference spectroscopy (RIfS). All of these methods sense changes of refractive index at an interface. RIfS relies on the shift of the interference pattern of white light reflected from a thin transparent film, which is caused by changes of the pathlength of the partial light beam traveling inside the transparent film for sensing purposes. It is the detec-tion method employed to develop a biomembrane sensor in this thesis and a detailed introduction to the technique will be given in chapter 3.
The field of label-free optical sensing has been dominated by SPR since the release of the first commercial instrument in the early 1990s. Surface plasmons are the particle equivalent of waves of electromagnetic radiation that can be formed under specific con-ditions at certain metal/dielectric interfaces. SPR is used to detect molecular binding events based on the behaviour of these surface plasmons. When gold- or silver-coated (typically glass) surfaces are exposed to monochromatic p-polarised light above the criti-cal angle of incidence, a sharp reduction (SPR minimum) in the amount of reflected light is observed due to the resonant transfer of the energy from the incoming light to surface plasmons generated at the metal/glass interface [10]. The specific angle (or wavelength) at which this occurs is extremely sensitive to the local optical properties of the interface.
Hence, the binding of molecules to the metal surface will alter the SPR minimum and can be used to detect molecular binding events [10]. In a typical experiment, one of the
molecular binding partners is coupled to the metal coated sensor surface, which is the key limitation of the technique. This coupling may affect the biological activity of the partner and further the resulting binding equilibria may be affected by mass transport effects localised to the sensor surface [16]. Current instruments can detect changes in mass∆m < 10 pg/mm2 on the sensor surface [10].
Optical waveguides have recently gained attention since they are compatible with existing SPR set-ups but offer higher sensitivity (∆m = (2.7 - 5) pg/mm2) [10]. Optical waveguides are specific structures that, when exposed to a wide spectrum light source, reflect light in a narrow band of wavelengths. The wavelengths of the reflected light are related to the materials that are used to form the waveguide, typically plastic (low refractive index) and a thin dielectric coating (high refractive index). Hence, in a manner analogous to SPR minimum changes, the peak wavelength value of the reflected light in waveguide-based technologies shifts to higher wavelengths in proportion to matter deposited on the sensor surface (increasing its dielectric permittivity) [17]. In addition, because waveguide-based surfaces may be generated with different physical composition, waveguide-based methods have broader application in that they may be used for cell-based assays as well, but the technique shares the disadvantages of SPR [10,18].
1.1.3. Acoustic Sensors
Another label-free sensing scheme is based on surface acoustic wave (SAW) devices.
SAWs generate and detect acoustic waves using interdigital transducers on the surface of a piezoelectric crystal [19]. In this way, the acoustic energy is strongly confined at the surface of the device in the range of the acoustic wavelength, regardless of the thickness of the complete substrate. For this reason, the wave is very sensitive towards any change on the surface, such as mass loading, viscosity and conductivity changes [19]. When immersed in aqueous liquids, SAW devices suffer from immense attenuation due to dis-placement of components perpendicular to the surface. The latter generate compression waves which radiate into the liquid and cause high attenuation of the device [20]. There-fore, research activity was initially focused on alternative acoustic wave types such as bulk acoustic waves (BAW) that mostly use thickness shear modes. These devices are commonly known as quartz crystal microbalances (QCM) [21]. A QCM detects mass loading or more precisely, the change in viscoelastic properties on an oscillating quartz through the change in frequency of the oscillation [22]. The quartz is set to oscillate via the reverse piezoelectric effect by applying a voltage to a gold electrode deposited on its surface, which is also where the interaction to be investigated takes place. Their resonance frequencies are usually in the range of 5 - 50 MHz. At higher frequencies the de-vices become too thin and thus too fragile for practical use. However, higher frequencies are most desirable, because the mass sensitivity increases with increasing frequency [23].
SAW-based biosensors allow the use of high frequencies in the range of several 100 MHz to GHz, implying higher mass sensitivities compared to QCMs.