ANALYSING THE BINDING OF CELL MEMBRANE BOUND MOLECULES

Field of the Invention

The invention relates to the field of analysing the binding of cell membrane bound molecules, in samples of whole cells, to specific binding molecules. The analysis may, for example, be an analysis as to whether the cell membrane bound molecules bind to the specific binding molecules or an analysis of (e.g. calculation of) kinetic and/or affinity parameters relating to the binding of the cell membrane bound molecules to the specific binding molecules.

Background to the Invention

The background to the invention will now be discussed with reference to the example application of calculating kinetic and/or affinity parameters concerning the binding of cell membrane proteins to specific binding molecules. Nevertheless, the invention is also applicable to other analyses of the binding of cell membrane bound molecules to specific binding molecules.

Within this specification and the appended claims the terms "cell membrane bound molecules" and "cell membrane proteins" includes molecules and proteins respectively which are bound to cell membranes directly or via other molecules. Cell membrane proteins are typically adhered to the cell membrane either by spanning the cell membrane and including a central hydrophobic region, or via a covalent bond to

one or more lipid molecules which are inserted into the cell membrane, or by adhesion (e.g. covalent attachment) to other cell membrane molecules.

Cell membrane proteins have great biological and biotechnological significance. Around 30% of all proteins encoded by the genome are cell membrane proteins. Cell membrane proteins mediate interactions between cells and extracellular components or other cells. Cell membrane proteins can also be important drug targets. Cells are in constant communication with their environment via cell membrane bound molecules. The binding of cell membrane molecules to specific ligands attached to the surface of other cells is pivotal in numerous physiological and developmental conditions such as leukocyte adhesion and rolling, and cell mediated immune reactions. Accordingly, it is desirable to be able to study the properties of membrane proteins whilst they remain attached to the cell membranes of whole cells. This enables properties of membrane protein interactions to be studied in their native environment, without requiring the proteins to be purified and separated from the membranes of whole cells.

Surface interactions involving cell-bound membrane proteins differ from the binding interactions which occur when soluble proteins bind immobilized ligands. This is because cell-bound membrane proteins are restricted in two dimensions due to spatial limitations imposed by the cell membrane. As a result, membrane-associated events are governed by two-dimensional (2D) kinetics and affinity, where association rate (k _a ) and binding affinity (K _A ) constants are typically expressed in units such as μm ² s ^"1 per molecule and μm ² per molecule respectively, instead of the traditional three-dimensional chemistry that occurs between soluble molecules, with k _a in units such as M ^"1 s ^"1 and K _A in M ^"1 . It is desirable to measure the two-dimensional kinetic and affinity parameters relating to the binding of cell membrane proteins to other molecules in order to better understand the binding mechanisms of molecular interactions between opposed membranes as these parameters are related directly to the function of membrane proteins.

It is known to measure two-dimensional kinetic and affinity parameters relating to cell membrane proteins using fluorescence and mechanical methods. However, multiple labour-intensive experiments are required to effectively measure 2D affinity and kinetics data.

A variety of biosensing techniques are known for measuring binding affinity and kinetic data relating to biological interactions between surface-bound molecules and water soluble or dispersed analytes. However, these techniques are not generally applicable to the measurement of binding affinity and kinetic data concerning interactions with whole cells and cell membrane bound receptors.

Optical biosensors, such as Surface Plasmon Resonance (SPR) devices, have not been effective due to the relatively large cell mass present in the sensor's evanescent field which produces a bulk response that is not sensitive to the number of membrane receptor/surface immobilized ligand complexes. In contrast, the present invention makes use of liquid-phase acoustic wave sensors which are sensitive to visco-elastic changes occurring close to the sensing surface, as well as mass coupling.

Accordingly, the invention aims to provide improved or alternative methods for analyzing the binding of cell membrane bound molecules to specific binding molecules whilst the cell membrane bound molecules remain bound to the membranes of whole, and preferably live, cells. Some embodiments of the invention seek to calculate two-dimensional kinetic and/or affinity parameters relating to the binding of cell membrane bound molecules, in samples of whole cells, to specific binding molecules.

Summary of the Invention

According to a first aspect of the present invention there is provided a method of analysing the binding of cell membrane bound molecules to specific binding molecules comprising the steps of:

(i) bringing a measurement sample of whole cells having cell membranes with target cell membrane bound molecules bound thereto into contact with the sensing surface of a liquid-phase acoustic wave sensor which generates an acoustic wave and produces a signal which is related to energy losses of the acoustic wave, which acoustic wave sensor has a sensing surface having specific binding molecules immobilised thereto which are operable to form specific bonds with target cell membrane bound molecules; and

(ii) monitoring the signal while specific bonds are formed between the specific binding molecules and the target cell membrane bound molecules of the measurement sample of whole cells.

The method may comprise the step of determining whether the cell membrane bound molecules bind to the specific binding molecules, using the monitored signal.

Alternatively, or additionally, the method may comprise the step of analysing one or more kinetic and/or affinity parameters relating to the cell membrane bound molecules within the measurement sample of whole cells, using the monitored signal.

The analysis of one or more kinetic and/or affinity parameters may be a calculation of one or more kinetic and/or affinity parameters relating to the binding of the cell membrane bound molecules within the measurement sample of whole cells to the specific binding molecules, using the monitored signal.

Preferably, at least one of the one or more kinetic and/or affinity parameters is a two- dimensional kinetic and/or affinity parameter.

The one or more kinetic and/or affinity parameters may be kinetic and/or affinity parameters of the binding of the cell membrane bound molecules to the specific binding molecules. For example, the one or more kinetic and/or affinity parameters may comprise or consist of the two-dimensional association rate (k _a ) of a cell membrane molecule bound with the specific binding molecules, the two-dimensional dissociation rate (k _d ) or the two-dimensional binding affinity (K _A ) of a cell membrane bound molecule with the specific binding molecules.

The method may comprise the step of determining properties of the cell membrane bound molecules which affect their interaction with the specific binding molecules. These properties can be deduced from the monitored signal. For example, the method may comprise the step of analysing (e.g. calculating) the rate of diffusion of the target cell-bound molecule within the cell membrane. (This two-dimensional rate of diffusion within a membrane should not be confused with the rate of diffusion in solution).

The method may comprise the step of determining whether further molecules compete with the specific binding molecules to bind the cell membrane bound

molecules. For example, the method may alternatively, or additionally, comprise the step of analysing (e.g. calculating) one or more kinetic and/or affinity parameters relating to the interaction between the cell membrane bound molecules within the measurement sample of whole cells and further molecules, using the monitored signal. The further molecules may, for example, be cell membrane bound molecules with which the target cell membrane bound molecules interact, or extracellular molecules which compete with the specific binding molecules to bind the target cell membrane bound molecules.

The step of calculating one or more kinetic and/or affinity parameters relating to cell membrane bound molecules using the monitored signal may optionally disregard the monitored signal for a period of time after cells were first brought into contact with the sensing surface. This enables changes in the monitored signal due to the diffusion of cells through the liquid to the sensing surface and their initial adherence to the sensing surface to be disregarded.

Preferably, the signal is monitored whilst changes in the monitored signal result predominantly (preferably substantially entirely) from the formation of bonds between the specific binding molecules and the target cell membrane bound molecules.

The step of calculating one or more kinetic and/or affinity parameter may comprise the step of determining the exponent, for a suitable base (such as 10 or e), of an exponential change in the monitored signal resulting from the binding of target cell membrane bound molecules to the specific binding molecules.

The analysis of the binding of cell membrane bound molecules to specific binding molecules, for example the analysis (typically calculation of) one or more kinetic and/or affinity parameters relating to the cell membrane bound molecules within the measurement sample of whole cells, may be an analysis of changes affecting the binding of cell membrane bound molecules to specific binding molecules.

The method may employ two different samples of whole cells each of which comprises target cell membrane bound molecules, wherein the samples of whole cells are different such that a different rate of binding of target cell membrane bound molecules to specific binding molecules arises when the two different samples of whole cells are introduced to equivalent sensing surfaces having specific binding molecules attached thereto.

Accordingly, the method may comprise the step of bringing a reference sample of whole cells having cell membranes with target cell membrane bound molecules bound thereto into contact with the sensing surface of a liquid-phase acoustic wave sensor which generates an acoustic wave and produces a signal which is related to energy losses of the acoustic wave, which acoustic wave sensor has a sensing surface comprising specific binding molecules which are operable to form specific bonds with target cell membrane bound molecules, and monitoring the signal while specific bonds are formed between the specific binding molecules and the target cell membrane bound molecules of the second sample of whole cells. The reference sample of whole cells may be used, for example, as a control or for calibration or comparative purposes, and may be brought into contact with the sensing surface of a liquid-phase acoustic wave sensor before, after, or simultaneously to the bringing of the measurement sample of whole cells into contact with the sensing surface of a liquid-phase acoustic wave sensor.

A plurality of signals measured while specific bonds are formed between specific binding molecules and target cell membrane bound molecules of samples of whole cells may be used to prepare calibration data for use when analysing the monitored signal while specific bonds are formed between the specific binding molecules and the target cell membrane bound molecules of the measurement sample of whole cells.

Thus, the method may be used to analyse the binding of cell membrane bound molecules in the reference sample of whole cells to specific binding molecules by comparing the monitored signal while specific bonds are formed between the specific binding molecules and the target cell membrane bound molecules of the reference sample of whole cells and the monitored signal while specific bonds are formed between the specific binding molecules and the target cell membrane bound molecules of the reference sample of whole cells.

The measurement and reference samples of whole cells may be brought into contact with the sensing surface of the same liquid-phase acoustic wave sensor or different liquid-phase acoustic wave sensors.

The method may be repeated using three or more than three samples of cells each of which comprises target cell membrane bound molecules, wherein the samples of whole cells are different such that a different rate of binding of target cell membrane

bound molecules to specific binding molecules arises when the different samples of whole cells are introduced to equivalent sensing surfaces having specific binding molecules attached thereto.

An array of liquid-phase acoustic wave sensors, each of which comprises an individual sensing surface, or a liquid-phase acoustic wave sensor comprising a plurality of separate detection regions and associated measurement channels for providing separate signals, within the same sensing surface using microfluidic devices, may be provided to analyse the binding of a plurality of samples (e.g. a plurality of measurement samples) of whole cells to specific binding molecules simultaneously.

The or each acoustic wave sensor may also be operable to carry out other measurements (i.e. non-acoustic measurements) to analyse the binding of whole cells to specific binding molecules. For example, the acoustic wave sensor may also be used as an optical sensor, for example, a Surface Plasmon Resonance sensor, or an electrochemical sensor.

Thus, the acoustic wave sensor may also be operable to carry out non-acoustic measurements (for example, optical or electrochemical measurements), and the method may comprise the step of monitoring an additional signal, derived from a non- acoustic measurement (for example, an optical or electrochemical measurement). Typically, the non-acoustic measurement (for example, the optical or electrochemical measurement) is a measurement indicative of one or more properties of matter on or adjacent to the sensing surface. The additional signal may be derived from a non- acoustic (for example, optical or electrochemical) measurement indicative of the formation of specific bonds with target cell membrane bound molecules. The additional signal may be derived from a non-acoustic (for example, optical or electrochemical) measurement which is sensitive to the amount of material (e.g. the number or mass of whole cells) on or adjacent to the sensing surface. The additional signal may be derived from a non-acoustic (for example, optical or electrochemical) measurement which is independent of the amount of material (e.g. the number or mass of whole cells) on or adjacent to the sensing surface. The monitored signal and the additional signal can be analysed to better determine one or more kinetic and/or affinity parameters relating to the cell membrane bound molecules within the measurement sample of whole cells.

The samples of whole cells (e.g. the measurement and reference samples of whole cells, or a plurality of measurement samples of whole cells, or a plurality of reference samples of whole cells) may differ in that they include different numbers of cells with the same average number of target cell membrane bound molecules. The samples of cells may differ in that they include cells with different numbers of target cell membrane bound molecules and/or different surface densities of target cell membrane bound molecules which are operable to bind the specific binding molecules. The samples of cells may have been subject to different chemical treatments which affect the number of target cell membrane bound molecules which are operable to bind the specific binding molecules. The samples of cells may have been cultured differently to affect the number of target cell membrane bound molecules which are operable to bind the specific binding molecules. The samples of cells may differ in that they include cells with target cell membrane bound molecules in different conformations with different affinities for the specific binding molecules.

Typically, the samples of cells will be selected such that in each case different surface concentrations of the target cell membrane bound molecules will be provided.

The samples of cells may differ in that they include cells which have been treated differently prior to binding to the specific binding molecules. This facilitates testing to establish whether a particular treatment, e.g. a particular cell growth regime or an interaction between a cell and a test agent leads to a change in the affinity and/or kinetics of a cell membrane bound molecule, for example because the growth regime or interaction affects the structure of the cell membrane bound molecule.

The samples of cells may differ in that they include different amounts of a reagent which affects the binding of target cell membrane bound molecules to specific binding molecules. The reagent may bind with the target cell membrane bound molecules. For example, the reagent may be another cell membrane bound molecule which binds with the target cell membrane bound molecules. The reagent may compete with the specific binding molecules to bind the target cell membrane bound molecules. The reagent may compete with the target cell membrane bound molecules to bind the specific binding molecules. The reagent may be operable to modify the target cell membrane bound molecules, for example, the target cell membrane bound molecules may comprise receptors and the reagent may be a ligand for the receptors.

The reagent may be a test agent and the method may be a method of determining whether the reagent interacts with the target cell membrane bound molecules or the specific binding molecules to cause a change in one or more kinetic and/or affinity parameters of the cell membrane bound molecule.

Preferably, the monitored signal is substantially independent of the mass within the penetration depth of the sensor. Preferably, the acoustic wave has a penetration depth of less than 500nm and more preferably less than 200nm from the sensing surface. More preferably still, the acoustic wave has a penetration depth of less than 100nm from the sensing surface. .Thus, the monitored signal should depend on the formation of specific bonds between cell membrane bound molecules and specific binding molecules and should not be affected by signal changes due to the presence of the cell mass in the sensing area The penetration depth is the distance from the sensing surface within which the amplitude of liquid oscillation decays to 1/e of the value of the amplitude of oscillation at the sensing surface.

The monitored signal will typically be related to the energy loss or dissipation of the acoustic wave which is generated, for example, the monitored signal may be related to the amplitude of the generated acoustic wave.

The liquid-phase acoustic wave sensor may be a Bulk Acoustic Wave type device, such as a Quartz Crystal Microbalance or Thickness Shear Mode resonator. In this case, the monitored signal will typically be related to the energy dissipation of the wave generated by the acoustic wave sensor.

The liquid-phase acoustic wave sensor may be an acoustic wave sensor which generates a shear wave; such Surface Acoustic Wave type devices can employ interdigitated transducers to generate a shear wave, such as a Love wave, Surface Skimming Bulk Wave, Acoustic Plate Mode, Bleustein-Gulyaev wave or Surface Transverse Wave. The monitored signal will typically be related to the energy loss or dissipation of the surface acoustic wave which is generated, for example, the monitored signal may be related to the amplitude of the generated surface acoustic wave.

The liquid-phase acoustic wave sensor may be an acoustic wave sensor using a thin membrane to excite an acoustic wave in a configuration known as Flexural Plate Wave or Lamb wave device.

The shear acoustic wave sensor may be a non-1 DT based device such as a device employing an electromagnetically excited shear acoustic wave.

The specific binding molecules are typically biological macromolecules, such as proteins, for example antibodies. However, the specific binding molecules may be other organic molecules, for example ligands for cell receptor proteins. The specific binding molecules may be immobilised on the sensing surface via intervening molecules, e.g. via a surfacing binding layer of molecules (such as a monolayer). The specific binding molecules may be bound to whole cells immobilised on the sensing surface.

The target cell membrane bound molecules are typically cell membrane molecules attached to the outside of cell membranes or located within cell membranes and operable to form bonds with extracellular specific binding molecules. The target cell membrane bound molecules may be cell membrane receptors and the specific binding molecules may be ligands for the cell membrane receptors.

Preferably, the cells are whole live cells.

According to a second aspect of the present invention there is provided a method of measuring the concentration of cell membrane bound molecules on a cell surface which are available to bind to surface-immobilised specific binding molecules, comprising the steps of:

(i) bringing at least one reference sample of whole cells having cell membranes with a known concentration of target cell membrane bound molecules bound to the surface thereof and available to bind to surface-immobilised specific binding molecules into contact with the sensing surface of a liquid-phase acoustic wave sensor which generates an acoustic wave and produces a signal which is related to energy losses of the acoustic wave, which acoustic wave sensor has a sensing surface having specific binding molecules immobilised thereto which are operable to form specific bonds with target cell membrane bound molecules, and monitoring the signal while specific bonds are formed between the specific binding molecules and the target cell membrane bound molecules of the reference sample of whole cells;

(ii) bringing a measurement sample of whole cells having cell membranes with an unknown concentration of target cell membrane bound molecules bound to the surface thereof and available to bind to surface-immobilised specific binding molecules into contact with the sensing surface of a liquid-phase acoustic wave sensor which generates an acoustic wave and produces a signal which is related to energy losses of the acoustic wave, which acoustic wave sensor has a sensing surface having specific binding molecules immobilised thereto which are operable to form specific bonds with target cell membrane bound molecules, and monitoring the signal while specific bonds are formed between the specific binding molecules and the target cell membrane bound molecules, if any, of the measurement sample of whole cells; and

(iii) determining the concentration of target cell membrane bound molecules which are available to bind the said immobilised specific binding molecules on the surface of the measurement sample of whole cells by comparing the monitored signals.

Typically, the step of comparing the monitored signals comprises comparing the rate of change of the monitored signals, optionally with normalisation. The monitored signals will typically be related to the energy loss or dissipation of the acoustic wave which is generated. For example, the monitored signals may be related to the amplitude of the generated acoustic wave. Thus, the step of comparing the monitored signals may comprise comparing the rate of amplitude change of the monitored signals normalised with respect to the absolute value of the amplitude change.

Preferably, step (i) is carried out in respect of a plurality of different reference samples of cells having cell membranes with a known concentration of target cell membrane bound molecules bound to the surface thereof and available to bind to surface-immobilised specific binding molecules and the plurality of resulting monitored signals are compared to determine the concentration of target cell membrane bound molecules and available to bind to surface-immobilised specific binding molecules. The plurality of resulting monitored signals may be used to provide a calibration curve.

The plurality of different reference samples of cells may have different concentrations of target cell membrane bound molecules bound to the surface thereof. The plurality of different reference samples of cells may comprise different concentrations of cells with the same known concentration of target cell membrane bound molecules bound to the surface thereof. The plurality of different reference samples of cells may have target cell membrane bound molecules with different availabilities to bind the said immobilised specific binding molecules, for example, because they have different configurations or have bound different concentrations of a competing molecule.

Preferably, the monitored signal is substantially independent of the mass within the penetration depth of the sensor, enabling different numbers of target cell membrane bound molecules to be bound to the specific binding molecules by varying the number of cells within the samples.

The method may comprise the step of measuring the number of cells which have adhered to the sensing surface to determine the number of target cell membrane bound molecules per cell.

The method may comprise the step of determining the availability of the target cell membrane bound molecules to bind the said immobilised specific binding molecules.

Further features of the second aspect of the invention correspond to those discussed in relation to the first aspect of the invention.

The invention also extends in a third aspect to a method of assessing the availability of the target cell bound molecules in a sample of whole cells to bind surface- immobilised specific binding molecules, measuring the concentration of cell membrane bound molecules on the surface of the treated cells which are available to bind to surface-immobilised specific binding molecules by a method according to the second aspect of the present invention, determining the concentration of cell membrane bound molecules on the surface of the treated cells and thereby determining the availability of the target cell membrane bound molecules of the treated sample of whole cells to bind the said immobilised specific binding molecules.

The method may comprise the step of treating a sample of whole cells prior to assessing the availability of the target cell bound molecules in the sample of treated

cells to bind to surface-immobilised specific binding molecules and thereby assessing the effect of the treatment on the availability of the target cell bound molecules in the sample of treated cells to bind to surface-immobilised specific binding molecules.

The treatment may comprise the addition of a test agent, to determine whether the test agent affects the availability of the target cell membrane bound molecules to bind the said immobilised specific binding molecules, for example by affecting the conformation of the immobilised specific binding molecules or by competing with the immobilised specific binding molecules to bind the target cell membrane bound molecules.

A method according to any one of the first three aspects of the invention may be part of a method of studying the interactions of cell membrane bound molecules and/or ligands for cell membrane bound molecules. A method according to any one of the first three aspects of the invention may be part of a method of studying interactions of immune system molecules. A method according to any one of the first three aspects of the invention may be part of a method of screening potential treatment agents. A method according to any one of the first three aspects of the invention may be part of a diagnostic method. A method according to any one of the first three aspects of the invention may be part of a method of screening surface structures which are candidates for use in synthetic transplants.

Brief Description of the Drawings

An example embodiment of the invention will now be illustrated with reference to the following Figures in which:

Figure 1 is a schematic perspective view of a Love wave sensor;

Figure 2 is a schematic cross-section through a Love wave sensor including a cell having cell membrane proteins;

Figure 3 is a graph of the change in amplitude (constituting the monitored signal) as a function of the number of bound cells during the binding of untreated, mild acid treated and from high density culture LG2 cells. Cell densities were calculated with an average accuracy of 10%;

Figure 4 is a graph of amplitude change as a function of time during the interaction of LG2 cells (2.5*10 ⁵ cells/ml) with the surface-immobilized anti-HLA antibody. Arrows (a) and (b) indicate the corresponding times for the addition of the cell suspension (10 μl min-1 ) and buffer (50 μl min-1 ), respectively. The binding curve is divided in two phases. The first one (~500 s after addition) represents the time required for cells to reach the surface and form initial tethers and is cell-diffusion limited. The second phase (shaded rectangle) depicts the formation of HLA/anti-HLA bonds;

Figure 5 is an optical microscopy photograph of LG2 cells on the biosensor surface after an acoustic experiment. Cytoplasmic protrusions are formed as a result of the formation of cell-surface contact points. (Scale bar: 15 μm);

Figure 6 is a graph showing real-time amplitude/time binding curves when various LG2 cell suspensions were added to immobilized anti-HLA antibody. Arrows (a) and (b) indicate the times for the addition of the cell suspension and buffer, respectively. In order from top to bottom the traces show the results for cell suspensions with concentrations of 6.0*10 ⁴ /ml, 1.0 ^χ 10 ⁵ /ml, 1.2 ^χ 10 ⁵ /ml, 1.5 ^χ 10 ⁵ /ml, 2.0 ^χ 10 ⁵ /ml, 2.4 ^χ 10 ⁵ /ml, 3.0 ^χ 10 ⁵ /ml, and 6.0*10 ⁵ /ml; and

Figure 7 is a plot of δ(dA/dt)/δA, where (dA/dt) is the rate of amplitude (A) change derived from real-time curves and δ(dA/dt)/δA the slope of (dA/dt) against A, versus the total number of cell-attached HLA-A2 molecules available for binding, i.e. C for the experiments illustrated in Figure 6. The two-dimensional rate constants k _a (in μm ² s ^"1 per molecule) and k _d (in s ^"1 ) can be derived from the slope and intercept with the ordinate, respectively, while the ratio of k _a /k _d is a measure of the two-dimensional binding affinity K _A (in μm ² per molecule).

Detailed Description of an Example Embodiment

Figure 1 is a schematic diagram of a Love wave sensor, shown generally as 1 , which functions as the liquid-phase acoustic wave sensor. The sensor comprises a quartz substrate 2, with interdigitated transducers (not shown) deposited thereon. A polymer guiding layer 4 is formed on the quartz substrate between the interdigitated transducers. A gold layer 6 is deposited on the polymer guiding layer to facilitate immobilisation of specific binding molecules. A Protein G layer 8 is adhered to the gold layer and antibodies 10, functioning as the specific binding molecules, are attached to the Protein G layer and thereby immobilised on the sensing surface

formed by the gold layer and immobilised proteins. Details of the construction of a Love wave sensor and formation of the antibody layer are described below.

In use, an oscillating electric potential is applied to the Love wave sensor via the IDTs to create a shear-horizontal surface acoustic wave. The phase and amplitude of the wave, which are related to the acoustic velocity and energy, respectively, are measured as a function of time through electrical connections to the IDTs and recorded for use in subsequent calculations. The polymer guiding layer serves as an acoustic waveguide by localizing the acoustic energy of the wave close to the sensing surface. All sensing occurs within the volume in contact with the device surface in which there is significant displacement as a result of the acoustic wave/liquid coupling. In the example embodiment, the acoustic wave sensor has a penetration depth of approximately 50 nm when in contact with pure water and increases with the square root of the liquid viscosity. The vast majority of the mass of cells which bind to the sensing surface is outside of the penetration depth.

With reference to Figure 2, whilst the phase and amplitude of the acoustic wave are monitored, whole live cells 12 including a target membrane protein 14 with which the immobilised antibodies can form specific bonds are brought into contact with the surface of the Love wave sensor. The approximate penetration depth of the acoustic wave sensor is shown by dashed line 16. The cells diffuse to the sensing surface and bind to the sensing surface by virtue of the specific binding interaction between the antibodies and the target membrane proteins. Specific bonds form between the antibodies and the target membrane proteins over a period of time (typically tens of minutes).

As we demonstrate below, the amplitude of the acoustic waves is correlated directly to the formation of specific bonds between the antibodies and the target membrane proteins and is not significantly affected by the mass of the cells. The amplitude is monitored while each of a plurality of samples of whole cells including different number of otherwise equivalent cells are brought into contact with sensing surfaces. Accordingly, by monitoring the amplitude of the acoustic wave while specific bonds are formed between antibodies and target membrane proteins, and comparing the change in amplitude with time for samples of whole cells with different numbers of cells, and therefore different maximum target membrane protein concentrations on the sensing surfaces, kinetic parameters concerning the interaction between the antibodies and the target membranes can be derived.

The rate of change of amplitude, δ(dA/dt), divided by the change of amplitude relative to a reference value, δA, is plotted against C, the two dimensional surface density of cell membrane HLA molecules. The 2-dimensional association and dissociation rate constants k _a and k _d are calculated from the slope of this plot and the intercept with the ordinate respectively. The two-dimensional binding affinity K _A can then be calculated as the ratio k _a /k _d .

Accordingly, this method allows the calculation of kinetic and affinity parameters of cell membrane bound molecules in situ whilst they remain bound to the membranes of live cells.

Experimental Evidence

In order to validate this approach, we carried out experiments which demonstrate that the change in energy loss (measured as amplitude in this example) with time in an arrangement such as that which is described in Figure 2 depends on the formation of specific antibody-membrane protein bonds rather than the mass of cells which is present.

The membrane receptor protein used was the HLA-A2 molecule, the most common class I Major Histocompatibility Complex allele in human populations, expressed in the B-lymphoblastoid LG2 cell line. The natural function of HLA-A2 is to present short endogenous peptides (8-1 1 residues) to the T cell receptors (TCRs) of T lymphocytes, which can trigger immune response. The specific binding molecules used in this example are surface immobilized anti-HLA-A2 monoclonal antibodies BB7.2; oriented on the surface through the F _c fragment (Saha et al., 2003; Fahnestock et al., S. R., 1986). The antibody is specific for the α chain (Parham and Brodsky, 1981 ) of HLA-A2 when the latter exists in a heterotrimer form consisting of α chain/β ₂ -microglobulin/peptide (Hogan and Brown, 1992). In this example, the heterotrimer form of the HLA-A2 molecules function as the target cell membrane bound molecules. Apart from the heterotrimeric form, HLA-A2 molecules on the cell membrane can be found in a heterodimeric form, i.e. just α-chain/β ₂ -microglobulin, or as single α-chains (Matko et al., 1994); α chains in the last two forms are not recognized by BB7.2.

Samples of cells containing different numbers of HLA-A2 heterotrimers were prepared using LG2 cells prepared under three different conditions: as untreated cells, as mild acid-treated and as untreated from high density cultures. Mild acid treatment removes bound peptides from the HLA groove (van der Burg et al., 1995), so treated cells have low heterotrimer numbers. In addition, cells grown in high density cultures will display high heterotrimer numbers on the cell surface (Matko et al., 1994). Using an indirect quantitative immunofluorescence assay and flow cytometry (described further below), the number of HLA-A2 heterotrimers (hereafter referred to as HLA or HLA-A2) on the cell surface was calculated. The values were: 3.7-5.7*10 ⁵ HLA-A2 molecules per untreated LG2 cell (12 experiments), 1-10*10 ⁴ molecules per mild-acid treated LG2 cell (6 experiments) and 9.7-10.0><10 ⁵ per LG2 cell from a high density culture (4 experiments).

Love-wave acoustic sensors are sensitive to mass loading and viscosity changes occurring within the sensor's penetration depth. Mass changes will only affect the phase of the acoustic wave while viscosity changes will affect primarily the amplitude. Upon addition of LG2 cells to the modified biosensor surface, the amplitude signal changed significantly and in proportion to the number of applied cells; in contrast, the phase signal displayed much lower sensitivity and could not discriminate between samples containing different number of cells. Closer inspection of the geometry of the biological sensing layer reveals that protein G would form a 5-10 nm thick layer while the IgG layer extends further from the gold surface for up to a further15 nm maximum (Harris et al., 1997). Upon cell addition, the HLA molecules on the cell membrane would interact with the immobilized IgG through the 7 nm long extracellular part of HLA-A2 molecules (Bjorkman et al., 1997). Taking into account cell membrane thickness (6-10 nm), it is clear that the vast majority of the cell mass will be outside the penetration depth (cell diameter 14.4±2.2 μm). For membrane HLA-A2, the calculated total mass-per-area values in the acoustic experiments are very low to cause a detectable phase change due to mass coupling. Instead, amplitude response is related to acoustic energy dissipated from the sensor surface to the liquid interface. Apparently, the cell membrane, in which the HLA molecules are embedded, acts as an effective damper by adsorbing energy through the HLA- A2/antibody bonds, hence the observed amplitude change.

In order to determine the exact effect of the number of cell surface HLA molecules on the signal change, experiments were performed using the three different types of LG2 cells, i.e. untreated, treated with a mild acid and from high density cultures. Cells

were injected at a low flow rate (10 μl min ^"1 ) to minimize shear stresses, at various concentrations, and the cell density on the biosensor surface was measured under the microscope after each experiment. It was found that, for the same number of cells bound to the device surface, the acoustic signal changed in proportion to the number of the HLA-A2 molecules on the cell membrane (Figure 3) proving that the formation of HLA/anti-HLA bonds is the major cause of the acoustic response and not the immobilization of cells on the sensor surface per se. Moreover, control experiments, which included the injection of HLA non-expressing K562 cells over the antibody-modified sensor surface and the addition of anti-HLA pretreated LG2 cells, also proved the specificity of the observed signal change to the formation of HLA/anti- HLA bonds.

Example Calculation of Two-Dimensional Kinetic Parameters and Affinity Constant

In order to calculate two-dimensional kinetic parameters and the affinity constant for the reaction between HLA-A2 and surface immobilized anti-HLA-A2 monoclonal antibodies, Love wave devices were prepared as described below. Anti-HLA-A2 monoclonal antibodies were immobilized on the sensor surface at a relatively high surface density (5.9±1.9*10 ³ molecules μm ^"2 ) to facilitate cell attachment.

Following harvesting, LG2 cells in PBS were injected over the antibody-functionalized sensor surface at a flow of 10 μl min ^"1 . Figure 4 shows a typical response of the acoustic signal amplitude change as a function of time. Before and after the injection of the cells, the acoustic signal was equilibrated with PBS as running buffer at a flow rate of 50 μl min ^"1 . Furthermore, to minimize any interference from diffusion-limited kinetics, only that part of the graph that fitted well (95%) to a single exponential curve, assuming a one-to-one reaction, was used in subsequent calculations, while the first 500 s following addition of cells were not included in the kinetic analysis. After this initial phase, signal changes observed should reflect the formation of new HLA/antibody bonds as a result of HLA-A2 lateral diffusion in the membrane followed by the formation of a larger cell-surface contact area (Figure 5).

The acoustic biosensor was found to be sensitive to the addition of as few as 6*10 ³ cells ml ^"1 or 156 cells immobilized on the sensor surface, which corresponds to a total mass of HLA molecules per sensor surface 0.6 pg mm ^"2 . This damping related detection is lower than the mass related detection of SPR for soluble molecules (100- 1000 pg mm ^"2 ).

Calculations

The amplitude signal change is a measure of the formation of HLA-IgG bonds, thus, kinetic information can be deduced from the real time binding curves for the cell- bound HLA/immobilized anti-HLA interaction. Experiments were carried out with cell suspensions containing different numbers of untreated LG2 cells (Figure 4a). For kinetic calculations, analysis previously described for 3D interactions (Saha et al., 2003; Karlsson et al., 1991 ) was applied according to the equation:

δ(dA/dt) / δA = k _a C + k _d (1 )

where (dA/dt) is the rate of amplitude (A) change derived from real-time graphs, C the concentration of the soluble analyte and k _a and k _d the association and dissociation rate constants, respectively. In our case, the three dimensional molar concentration C is replaced by the two dimensional HLA-A2 surface density C ^2D , which should reflect the number of cell membrane HLA molecules available for binding, and k _a , k _d are the corresponding 2D rate constants.

The number of laterally mobile HLA molecules available for interaction on each cell surface is given by (Zhu et al., 2007):

[HLA] = N _HLλ x (f / Seen) (2)

N _HLA being the total number of HLA molecules on the cell surface, f their fractional mobility, and S _ce ιι the cell surface. The surface area (S _ce ιι)of the cells was calculated to be 633.15 μm ² (S _ce ιι = πd ² , where the cell's diameter d was measured from microscopy photos to be 14.4 μm). In order to account for the irregularities of the cell membrane (Zhu et al., 2007), S _ce ιι was multiplied by the surface roughness factor of 1.8, resulting in a final value of 1 172.6 μm ² . Based on the average number of 4.7x10 ⁵ HLA-A2 receptors per untreated cells and the receptors' fractional mobility f (Bierer et al., 1987) of 93.5%, [HLA] is calculated to be 374±80 molecules μm ^"2 .

In our experiments, the total number of HLA receptors per cell remains constant and the only variable is the total number of applied cells. This implies that 374 molecules μm ^"2 is the maximum [HLA] that can be presented to the sensor's surface and will correspond to the maximum coverage of the surface by cells. In practice, the

maximum coverage was observed at cell suspensions equal or higher to 6x10 ⁵ cells/ml (C _max ), as detected acoustically through δA measurements. For cell suspensions lower than C _max , the HLA surface density per cell will still remain 374 molecules μm ^"2 ; however, the effective [HLA] sensed by the sensor will be a function of the coverage of the sensor surface by the cells, i.e. of d ^O6 " ⁵ /d _max ⁰⁶ " ⁵ , where d is the cells' surface density on the device. The observed linear relationship between C ^cells and d ^cells up to the saturation level allows rewriting the above ratio as C ₁ ⁰⁶¹¹⁵ ZC _m3x ⁰⁶ " ⁵ . Given the above analysis, the C ^2D , which is the equivalent to the soluble analyte concentration used in equation (1 ), can be calculated from:

C ^2D = [HLA] ^χ C ⁰⁶ " ⁵ /C _max ⁰⁶ " ⁵ (3)

Plotting δ(dA/dt)/δA versus C ^2D allows for the measurement of the 2D association and dissociation rate constants: k _a from the slope (4.88*10 ^"6 μm ² s ^"1 per molecule) and k _d from the intercept with the ordinate (2.07x10 ^"5 s ^"1 ). The two-dimensional binding affinity K _A can then be calculated: K _A = k _a /k _d = 0.236 μm ² per molecule. These binding parameters relate directly to the native function of cell membrane molecules.

The above analysis was performed given a constant surface density for immobilized IgG. Even though dissociation was not observed for whole cells, it can occur for single HLA/anti-HLA bonds. Rebinding of the dissociated HLA can readily occur, given the relatively high surface antibody density (5.9±1.9*10 ³ molecules μm ^"2 ). Moreover, the lateral diffusion of HLA molecules on the membrane of B- lymphoblastoid cells is higher (Bierer et al., 1987) (1.75*10 ^"1 μm ² s ^"1 ) than the measured association rate constant for bond formation (4.88* 10 ^"6 μm ² s ^"1 ); hence the rate constant is the limiting step and was, therefore, what was measured.

Conclusions

Accordingly, we have demonstrated that an acoustic wave sensor can be employed for assaying the binding of receptor-bearing cells, i.e. leukemic cells expressing class I Major Histocompatibility Complex (MHC) molecules, to a specific monoclonal antibody immobilized on the sensor surface. The measured signal (amplitude) changed as a function of time upon addition of viable cells to the modified sensor surface. The time course of the change in the measured signal was found to correlate directly to the number of cell-membrane molecules specifically attached to

the immobilized antibodies. This finding allowed for the calculation of two- dimensional kinetics and affinity parameters.

The sensitivity of acoustic damping to the number of cell/surface specific bonds provides a unique sensing mechanism for investigating membrane interactions. Clearly, the proposed label-free and non invasive acoustic biosensor could be further applied to characterize various membrane-associated events. Examples of interactions which could be studied include T cell receptor/MHC and MHC/antigenic peptide interactions using whole cells, thereby avoiding the need for membrane protein purification and/or reconstitution.

Alternative Application

We have also demonstrated that δ(dA/dt)/δA is proportional to C, the total surface density of cell bound molecules which are available to interact with an immobilised ligand. Thus, the total surface density of cell bound molecules which are available to interact with an immobilised ligand within a measurement sample of whole cells having an unknown total surface density of cell bound molecules which are available to interact with an immobilised ligand can be determined from a calibration curve of δ(dA/dt)/δA. A calibration curve of δ(dA/dt)/δA can be determined either by using reference samples having different concentrations of cells, each of which has the same the total surface density of cell bound molecules which are available to interact with an immobilised ligand, or by using different reference samples of cells having different total surface densities of cell bound molecules which are available to interact with an immobilised ligand.

Because δ(dA/dt)/δA is a function of the total surface density of cell bound molecules which are available to interact with an immobilised ligand, a comparison of δ(dA/dt)/δA between two samples with the same total surface density of cell bound molecules but which have been treated differently (for example, by exposing one of them to a test agent) enables changes in cell bound molecules which affect their availability to interact with an immobilised ligand, e.g. conformation changes or the binding of competing molecules, to be detected or quantified.

Further Example Applications

The invention can be applied to the study of cell membrane molecules and their ligands. For example, the invention can be employed to calculate the two- dimensional kinetics and affinity of various molecular interactions that mediate physiological and developmental conditions. One specific example is the study of the interactions of leukocyte homing receptors. These receptors determine the arrest and extravasation of neutrophils and other leukocytes in response to infection. In an example, leukocytes carrying the receptors are introduced to vicinity of the sensor surface where the ligands would be immobilized. This setup resembles the physiological condition as the ligands are normally expressed on the surface of the endothelial cells that form blood vessels. Various pumping rates can be employed to resemble the physiological blood flows. This example can be extended to the study of receptors that mediate tumor metastasis, since the mechanism of action is similar.

Another example application is the study of immunologically important interactions. Such interactions include T cell receptor/major histocompatibility complex molecules, NK receptors and their ligands and Toll-like receptors and their ligands. These interactions have been extensively studied using soluble molecules. However, in order to better understand the onset and progress of immune responses, it would be preferable for these interactions to be studied in a more physiologically relevant context, i.e. the cell membrane and using whole cells. This can also provide insight into the events following molecular interaction. In an example, immobilized ligands are provided on the sensor surface and cells carrying the receptors are introduced to the vicinity of the sensor surface. The real-time monitoring of the bond formation between cell and substrate provides information on the extent of the interaction. Besides the kinetics and affinity for the molecular interaction, it is important to investigate the downstream events, i.e. whether the immune cells used have been activated. The mechanism of activation can be investigated using various ligands at the sensor surface or a different extent of immobilization or even a different type of immobilization (i.e. insertion of ligands into supported lipid bilayers in order to be laterally mobile).

The invention can be also applied to the screening of pharmacological agents. One particular field that can benefit is the screening and analysis of therapeutic antibodies. The invention can be used to investigate the two-dimensional kinetics and affinity for the interactions of therapeutic antibodies and their molecular targets on the surface of

particular cell types, i.e. tumor cells. The experimental procedure can comprise immobilizing the antibody of interest on the sensor surface and then adding the cell type that presents the molecular target of the antibody at its surface. In that way, the efficacy of the therapeutic antibody can be easily evaluated. Another example of application of the invention to drug screening is the field of viral attachment and entry. Viruses specifically attach to and enter in certain cell types. They do that via their viral receptors that recognize and bind their ligands on the surface of these cell types. The invention can be employed to study the molecular interaction events of viral attachment. In this paradigm, the viral receptors can be immobilized on the sensor surface and various cell types applied over the surface. This experimental setup can yield the cell type specificity of viruses as well as the kinetics and affinity of the molecular interaction of viral receptors and their ligands. Moreover, using this setup, one can test for pharmacological factors that may inhibit viral attachment and entry. Potential inhibitors can be added in real-time during the monitoring of the interaction of viral receptors and their target cells. The successful inhibitors will result in lowering the monitored signal change since no binding would occur. Besides being added in real-time, potential inhibitors can be used to treat the target cells prior to addition over the sensor surface.

Diagnostics is a another field which could benefit from the current invention. The invention can be applied to detect various types of molecules or cells in blood or other biological fluids. Certain molecules (i.e. antibodies, specific receptors/ligands) can be immobilized to the sensor surface to form a biorecognition surface for biomarkers of disease, cancer or viruses. It can also offer detection of pathological cells, such as leukemic cells or virus-infected cells. Another example of the use of the invention in diagnostics can be immunophenotyping, i.e. the identification of antigens of blood groups or the major histocompatibilty complex with the use of antibodies immobilized on the sensor surface. This can become a fast screening assay of blood cells for the antigens that determine blood transfusions and organ transplantations.

The invention can be also applied to the screening of biocompatible surfaces that are used in synthetic transplants. The working example can be to prepare sensor surface as in the transplants and then apply certain cell types. This will allow checking for the way and extent cells interact with the transplant surface, a factor that usually determines the efficacy of biocompatible surfaces. Moreover, certain molecules are usually attached on the surface of transplants. Such molecules, which can be certain receptors or growth factors, are usually the bio-active part of the transplant. The

invention can be used to measure interactions between the immobilized molecules and their target cells. This can yield the kinetics and affinity of the molecular interaction of the immobilized molecules and their ligands in an environment close to the physiological conditions, i.e. the cell surface.

Materials and Methods

Surface Acoustic Wave Sensor Fabrication

1 10 MHz quartz devices were fabricated on 0.5-mm thick Y-cut piezoelectric quartz crystals. The interdigitated transducers, composed of a 210-nm thick Cr/Au (10/200 nm) electrode, consisted of 80 pairs of split fingers with a periodicity of 45 μm. The devices were coated with a 0.7 μm thick poly(methylmethacrylate) (PMMA) (Aldrich) layer on top of which 20 nm of gold were sputtered using a BAL-TEC SCD 050 sputter coater. The acoustic devices were mounted on a special holder and liquid was pumped through on the area between the IDTs using a peristaltic pump (Gilson) and a flow-through cell. The flow cell was sealed on the surface by using a custom- made rubber gasket exposing a sensing area of 12 mm ² . A Hewlett-Packard 4195A network analyzer was used to monitor the amplitude and phase of the wave and LabVIEW (National Instruments) software for collecting acoustic data. Experiments were run at least in triplicates. Following the end of acoustic experiments, the biosensor surfaces were observed under a Nikon Eclipse E800 microscope and photographs were taken with an attached Nikon Coolpix E5400 camera. Cells on the sensor surface were counted from at least three different areas.

Further details of the construction of the acoustic waveguide described herein have been published in Gizeli et al., 1992; Gizeli et al., 1997; Gizeli et al., 2003 and Saha et al., 2003.

Preparation of Sensing Surface

Freshly plasma-etched gold surfaces were incubated with a protein G (Calbiochem) solution (1 mg ml ^"1 ) which was left to adsorb for 1 h at room temperature. Following protein adsorption, the devices were inserted in the device holder, washed and left to equilibrate with PBS (Sigma) at a flow rate of 50 μl min ^"1 . Antibody solutions of 0.1 , 0.5, 1 or 10 μg ml ^"1 were pumped over the biosensor surfaces under 25 μl min ^"1 flow rate and the interaction with the gold-adsorbed protein G (functioning as the specific

binding molecules) was monitored in real time. IgG surface densities were measured from surface plasmon resonance (SPR) experiments (SR7000, Reichert) by injecting the antibody over a protein G-modified sensor surface.

Cell Cultures and Treatments

The EBV-transformed human B-lymphoblastoid cell line LG2 (HLA-A ^* O2O1 Y) and the chronic myelogenous leukemic K562 cells (HLA-A ^" B ^" C ^" ) were used in this work. RPMI 1640 (GIBCO Inc.) supplemented with 1 mg L ^"1 gentamycin and 10% of fetal bovine serum was used as culture medium. Culture flasks were kept in humidified 5% CO ₂ atmosphere at 37° C. Medium was exchanged every 2-3 days. The cell density was 3-8*10 ⁵ cells ml ^"1 in normal cultures and 2-3*10 ⁶ cells ml ^"1 in high density cultures. Viability was checked prior to experiments by the trypan blue exclusion technique and cell number was counted on a Neubauer slide. Dead cells were never over 5% of the total. Cells were washed with PBS (Sigma), centrifuged at 250 g and resuspended in PBS prior to addition to the sensor surface. In order to remove cell surface HLA-A2 associated peptides, LG2 cells were briefly (90 s) treated with ice-cold pH 3.2 citric acid-Na ₂ HPO ₄ buffer (mixture of an equal volume of 0.263 M citric acid and 0.123 M Na ₂ HPO ₄ ) (van der Burg et al., 1995) and then washed with PBS.

Flow Cytometry

In order to measure the number of cell surface HLA-A2 molecules, cells were incubated with the anti-HLA-A2 monoclonal antibody BB7.2 (Becton Dickinson) and a FITC-conjugated anti-mouse IgG secondary antibody (Dako), both at saturating conditions. Samples were analyzed in a FACS scanner (FACS Culiber, Becton Dickinson). Light scattering and FITC fluorescence data were collected on 4-decade logarithmic scales. The Ql Fl KIT ^® (Dako) was used to perform a quantitative immunofluorescence indirect assay using different populations of plastic beads with a known specific number of attached mouse IgG molecules. The number of cell surface HLA-A2 molecules was calculated according to manufacturer's instructions.

The following documents referred to in this description are incorporated herein by virtue of this reference:

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Further modifications and variations may be made within the scope of the invention herein disclosed.