Electrically excitable cells, such as cardiomyocytes and neurons, have been shown to produce fast optical signals that are the result of light scattering and birefringence changes associated with membrane depolarization [ 1–3 ]. In order to develop an understanding of how a population of cells are organized, exchange, and process information a new sensing technology is required, one that has single cell resolution over a relatively large network. Surface plasmon resonance (SPR) sensors possess highly sensitive resonance conditions, which make them capable of detecting variations in refractive index with a high spatiotemporal resolution, label-free thus enabling long-term recording.

SPR occurs when p-polarized light incident on a noble metal, at a speci c angle of incidence, couples to the free electrons in the metal resulting in all the incident energy generating surface plasmons (SPs) [ 4 ]. At the angle of incidence where the plasmon coupling occurs, the intensity of the re ected light drops and an evanescent wave is generated in both the metal and dielectric ( gure 1(a)). This evanescent wave penetrates into the dielectric to a depth of about 150 nm, depending on the properties of the metal used [ 5, 6 ]. SPR is very sensitive to perturbations of refractive index within this evanescent eld; therefore, when there is a change in the refractive index of the dielectric medium, the characteristics of the light wave coupled to the surface plasmon changes and the resonance conditions are altered ( gure 1(b)). This allows SPR to be used for label-free imaging of both structural (sensitive to local variations in density) [ 7–10 ], and functional (sensitive to movements of matter) [ 11–14 ] features.

SPs cannot be excited by the incident light directly. At a given photon energy (hω) the wave-vector of the incident p-polarized light must be increased so the photons can be coupled into plasmons [ 15 ]. Passing the illumination light through a high refractive index dielectric material, such as glass, can shift the wave-vector of the incident light enough to excite SPs. Excitation methods based on attenuated total re ection (ATR) were demonstrated in the late sixties by Otto [ 16 ] and Kretschmann and Raether [ 17 ]. The Kretschmann ATR conguration is one of the most popular methods for exciting SPs and has a thin metal lm deposited on top of a prism surface.

The Kretschmann–Raether con guration allows a large eld of view (FOV) (up to cm). However, this con guration is not well suited to imaging because it requires a very shallow angle of incident light. The physical constraint of the prism limits the numerical aperture (NA) and magni cation of the imaging system resulting in the con guration having a reduced spatial resolution [ 18 ]. Additionally, the FOV is angled relative to the objective resulting in an anisotropic distortion. Using a high NA objective lens instead of a prism to shift the wave-vector of the incident light resolves the issues of spatial distortion by keeping the object plane parallel to the image plane [ 18, 19 ]. Using a high NA objective also allows higher magni cation and, because of the high NA, the spatial resolution is much improved and can be diffraction limited (~300 nm).

Despite these caveats, using SPR for functional imaging by detecting refractive index variations within the evanescent (a) (b)

eld is possible. To detect variations in the refractive index of the dielectric medium interfaced to the metal surface, a number of detection schemes can be employed: (i) by tracking the angle of minimum light intensity (angular modulation) [ 20 ]; (ii) by monitoring the intensity with the angle of incidence xed at the position with the steepest gradient on the SPR curve (intensity modulation) [ 21, 22 ] or using differential intensity detection approaches [ 23, 24 ]; (iii) by detecting the phase of the re ected light [ 25 ]. The detection of small refractive index changes over a relatively large volume has been successful on some sensors based on an intensity modulation scheme down to a sensitivity (Δnmin) of 10−6 refractive index units (RIUs) [ 21, 22 ]. Better sensitivity levels have been achieved using other detection methods down to 10−7 RIUs using angular modulation [ 20 ], however, it is experimentally more complicated to monitor the angle of minimum re ection at fast sampling rates. A trade-off between sensitivity and sampling rate will be necessary if SPR is to be used for fast, functional imaging.

The spatial resolution of the SPR imaging system is limited in the direction perpendicular to the propagation of the SP [ 26 ]. Therefore, complementary information can be obtained from other methods including total internal re ection microscopy (TIRM). TIRM sees an evanescent wave generated at an interface separating two media with different refractive indices when the incident light is totally internally re ected (see supplementary information (stacks.iop.org/ JPhysD/52/104001/mmedia)) for discussion on total internal re ection). Discontinuities in the refractive index of the cell membrane give rise to differences in the re ected light intensity. TIRM has been used to study cell-surface interactions for improved drug delivery systems [ 27 ].

The evanescent wave generated at the interface is capable of exciting uorophores, and since the intensity of the evanescent eld decays exponentially with distance from the interface, only uorophores within ~100 nm are excited. In comparison with standard epi- uorescence the background is 2000 times lower because no out of focus uorescence gets excited thus resulting in a high signal-to-noise ratio (SNR) [ 28 ]. Depending on the substrate this technique is known as total internal re ection uorescence microscopy (TIRF) [ 29, 30 ] or SPR uorescence microscopy (SPRF). TIRFM has not been demonstrated in this study, however, we have shown the capability of the imaging system to excite uorescence using SPR. SPRF takes advantage of the enhanced electromagnetic eld intensity that occurs when plasmons are excited, resulting in ampli ed uorescence signals (see [ 31 ] for a detailed review).

Applying SPR imaging to research questions in cell physiology and pharmacology requires the development of a multimodal system where complementary information can be obtained. Information acquired from label-free techniques, such as SPR and TIRM can be validated with well established uorescent labelling methods. Likewise, the system can be used to validate the functional imaging capabilities of SPR.

In this paper, we present a multimodal imaging platform that includes SPR microscopy using a high NA objective applied to live cell imaging with the following capabilities: (i) the SPR system features a wide FOV providing the ability to study ~40 cells simultaneously, with subcellular resolution. (ii) The SPR system is used to image neuronal cells while resolving axons and dendrites. We comment on the factors that affect the resolution of ne neuronal processes. (iii) We show that complementary information on the imaging resolution can be obtained from TIRM and demonstrate a combination of microscopy systems including bright- eld, epi- uorescence and SPRF that can be applied to cell physiology. (iv) We combine a number of the imaging techniques to demonstrate a typical experiment where structural and functional imaging is desired and in doing so we verify the ability of the system to study spatiotemporal cellular functions by imaging localized contractions of stem cell-derived cardiomyocytes. (v) We describe a detailed design of this platform to enable ease of implementation and characterization for use by the cell physiology community.

2.1. Optical system design The custom-built multimodal imaging system we have developed combines SPR imaging with a number of microscopy sub-systems that include bright- eld, epi- uorescence, TIRM and SPRF. Figure 2 shows a schematic of the optical system and optical pathways. The system is also capable of performing electrophysiological measurements simultaneously using microelectrode technology. The rig was mounted on a oating optical table  (Thorlabs, Newton, NJ, USA) to minimize mechanical perturbations, to which such sensitive methods are highly susceptible.

A 640 nm light emitting diode (LED) was used in conjunction with Köhler illumination to illuminate the back focal plane (BFP) of the Nikon Oil Objective lens (CFI Apo TIRF 60×, NA = 1.49, oil-immersion lens, Nikon Inc., Tokyo, Japan) with a uniform light intensity [ 32 ]. The BFP was monitored on a complementaty metal oxide semiconductor (CMOS) camera (Unibrain, 640 × 480 pixels, 5.6 μm pixel size, Fire-i™, Unibrain Inc., San Ramon, CA, USA), shown in gure 3(a).

Re ection microscopy was performed with the 470 nm LED uniformly illuminating the BFP of a water-dipping objective (60×, NA = 1.00, Nikon Inc, Tokyo, Japan). The re ected light passed through the same objective lens, and was imaged on to a CCD camera (Grasshopper3, 1920 × 1440 pixels, 4.54 μm pixel size, S3-U3-28S5M-C Point Grey, Richmond, BC, Canada). Transmission microscopy was performed using the 470 nm LED and the transmitted light was imaged onto a second CMOS camera (640 × 480 pixels, 9.9 μm pixel size, SV643M, EPIX Inc., Buffalo Grove, IL, USA). The 470 nm and 590 nm LEDs could also be used for exciting uorescence, in combination with appropriate emission lters.

TIRM was achieved using the 640 nm LED, in conjunction with Köhler illumination to uniformly illuminate the BFP of the objective. The aperture diaphragm (D1) within the Köhler illumination was closed so that contrast was obtained in re ection by frustration of the critical angle (see supplementary information for information on the critical angle). The re ected light was imaged on the Unibrain camera.

When the objective pupil is lled with plane-polarized light from the 640 nm LED ( gures 3(a) and (b)), in the sector where the azimuthal angle is 0° the polarization state is pure p-polarized light, while pure s-polarized light arrives at the sample where the azimuthal angle is ±90°. The dark arcs visible in gure 3(a) show the angle of incidence where the light is coupled into SPs and is therefore no longer being re ected. There is no dark band in the vertical direction (±90°) due to the light being pure s-polarized, and therefore incapable of exciting SPs.

SPRF was performed by diffusional loading of Alexa Fluor®, 680-dextran (3kDa, Invitrogen™, D34681) into the cytoplasm of the cell of interest, using a micropipette attached to the cell soma. The uorescent dye was then excited using the 680 nm superluminescent LED (SLED, 1 mW, Superlum Diodes Ltd, Cork, Ireland) which was tuned to excite plasmons beneath the cell of interest. The SPR uorescence was imaged on the Point-grey Grasshopper3 camera. SPRF microscopy could be performed to con rm that the cells on the sample were in the evanescent eld prior to functional SPR imaging.

The SPR imaging system was developed around a high NA (NA = 1.49) objective lens and is able to exploit both angular and intensity modulation by either shifting the illumination angle (laterally translating the focal spot in the BFP) or holding this angle constant at the angle of maximum SPR gradient and monitoring the resulting intensity.

Angular modulation uses monochromatic light to excite the SPs, and the excitation can be seen as a dip in the angular spectrum. The detector senses a shift in the angle where the re ected light intensity is at a minimum. Angular modulation was achieved using the 680 nm SLED, generating light that is focused onto the objective BFP, resulting in an illuminating beam at the sample of an adjustable, narrow range of angles. A SLED was used over a laser or LED because they have a high spatial coherence but a low temporal coherence, allowing the light to be focused to a very tight spot [ 33, 34 ], without suffering from the effects of laser speckling [ 35 ]. Additionally, SLEDs have extremely low noise [ 36 ], which is advantageous in the high-speed acquisition of rapid signals such as electrical activity in cells.

The output of the ber-coupled 680 nm SLED was collimated using an achromatic lens (f = 100 mm) and passed through a polarizer so that only p-polarized light was incident on the sample. The polarized light was focused to a diffraction-limited spot on the BFP of the objective lens with a second achromat lens (f = 75 mm). The diffraction-limited spot, along with the polarizer, reduces the level of background caused by the s-polarized light that is unable to excite SPs. The radial position of the illumination spot in the BFP dictates the angle of incidence at the sample. The light scattered off refractive index discontinuities in the sample when cells are cultured on the gold coverslips returns from the same elevation angle but all azimuthal angles. This illustrates that momentum is conserved during plasmon scattering, as shown in gure 3(d). To change the angle of incidence at the sample the focused SLED beam was scanned across the BFP of the sample using a stepper motor traveling at 0.1 mm s−1. The re ected light was imaged on the EPIX camera at 24 fps, with an exposure time of 0.124 ms to provide re ection values as a function of the angle of incidence. Regions of interest were selected on the surfaces using Image-J and the z-axis pro le was exported to MATLAB ( gure 3(c)).

SPR sensing with intensity modulation (IM) works by xing both the angle of incidence and wavelength and measuring the strength of the coupling between the light wave and the SP; detection is achieved by measuring the change in the intensity of the re ected light ( gure 1(b)). Therefore, for IM detection the angle of incidence (θ) was xed at the angle with the steepest gradient, ΔR/Δθ (around 30% of the minimum intensity recorded at the trough of the SPR dip, marked on gure  3(c)) and the variation of the re ected light intensity was monitored on the EPIX camera. 2.2. SPR sensor preparation The glass coverslips (19 mm, Karl Hecht GmbH and Co KG, Sondheim, Germany) were treated with (3-mercaptopropyl) tri-methoxysilane (MPTS, Sigma Aldrich, 174617) to present thiol groups prior to thermal metal evaporation [ 37 ]. MPTS was used instead of the usual chromium or titanium adhesion layer to preserve the surface plasmon quality [ 38 ]. Silanisation was performed immediately after solvent and plasma cleaning. The plasma cleaned coverslips were immersed in a boiling solution of 50:1:1 isopropanol: MPTS: deionised water for 10 min, then rinsed with isopropanol, dried under nitrogen and transferred to an oven at 100 °C for 10 min. The boiling, rinsing and baking steps were repeated three times.

The MPTS glass coverslips were coated with ~50 nm thick gold by thermal metal evaporation. Gold evaporation was carried out using a thermal evaporator (E306A, Edwards, Burgess Hill, UK). Gold wire was placed in a lament and the chamber was pumped down to 10−1 mbar with a rotary pump, then to 10−7 mbar with an oil diffusion pump, while being cooled with liquid nitrogen. A current of ~30 A was passed through the lament, so the metal melted and evaporated onto the glass coverslips held above the source with rare earth magnets. An Intellemetrics IL100 quartz crystal microbalance was used to estimate the evaporation rate and lm thickness. The thickness of the gold coverslips was measured with a spectroscopic ellipsometer (Alpha-SE, J.A. Woollam, Lincoln, NE, USA) at a xed incident angle (70°) and with a wavelength range between 380 to 900 nm. (a) (b) (c) (d) The sensitivity of the planar gold surfaces was experimentally determined using sodium chloride (NaCl) salt-water solutions. Different concentrations of salt solutions were manually added to a chamber on top of a gold-coated coverslip using a 1 ml pipette whilst the re ected SPR light intensity response was measured. Deionized water was added as the

rst sample to nd the re ection gradient maximum and the SLED was positioned here for the entirety of the experiment. While keeping all other parameters xed, the refractive index of the media was adjusted using different concentrations of salt solution. The change in refractive index was calibrated with an Abbe Refractometer (Abbemat 200, Anton Paar, Graz, Austria). The re ected light from the sample plane was recorded using the EPIX camera (12-bit) at ~286 Hz. To allow the solution to stabilize, 30–60 s was left between each step. The resulting intensities were exported using Image-J and further analysis was performed in MATLAB. 2.4. Substrate preparation For cell-based experiments on glass, the coverslips were treated with poly-L-lysine (PLL, Sigma Aldrich, P4707). The PLL coated glass coverslips were prepared by immersion in PLL solution for 5 min followed by thorough rinsing with deionized water.

For cell-based experiments on planar gold, the coverslips were treated with 11-Amino-1-undecanethiol (AUT, Sigma Aldrich, 674397), to generate a self-assembled monolayer [ 39 ]. First, the gold coverslips were solvent cleaned, followed by an oxygen plasma treatment.

The AUT was dissolved in ACS grade Ethanol at 1 mM, which was then poured into an immersion cylinder to completely cover the gold. Immersion time was not less than 18 h. Following immersion, the solution was drained from the cylinder and the gold was thoroughly rinsed, rst with ethanol to remove the bulk AUT solution, and then with distilled water to remove the cytotoxic ethanol. The coverslips were dried thoroughly with N2 before cell plating. 2.5. Cell culture The 3T3 broblast cell line was maintained in Dulbecco’s Modi ed Eagle Medium (DMEM, Gibco, 11960) with 10% fetal bovine serum (FBS, Gibco, 10270). Cells were subcultured and seeded on to AUT treated gold coverslips at a density of 0.3 × 106 cells.

The stem cell-derived cardiomyocytes were differentiated on AUT treated gold coverslips from human embryonic stem cells (hESC-CM) following the protocol in [ 40 ].

Cultures of primary hippocampal neurons were dissected from E18 Wistar rat embryos following a standard protocol [ 41 ]. The primary rat hippocampal neurons were plated on the functionalized coverslips with ~150 000 dissociated cells in 500 μl media. 3. Structural imaging 3.1. Optical system characterisation 3.1.1. Field of view (FOV). Functional imaging of excitable cells using SPR microscopy requires performing measurements from a number of single cells to inspect the cell–cell interactions and to replicate single cell measurements. Therefore, the system has been designed with a wide-FOV to allow imaging of multiple single cells with subcellular resolution. This has been achieved by deliberately reducing the magni cation of the imaging system.

The FOV was increased by mismatching the objective and tube lens to reduce the effective magni cation of the 60× oilobjective. The Nikon objective lens is designed for a tube lens with a focal length of 200 nm, however, we used a 60 mm tube lens. This reduces the effective magni cation to 18× and enlarges the FOV to up to 500 μm, depending on the size of the camera sensor.

The number of pixels in the EPIX camera sensor is 640 × 480, at 9.9 × 9.9 μm each, so the FOV has been increased to 350 × 250 μm from 105 × 80 μm. However, the back aperture of the objective lens restricts the usable FOV to a circular diameter of 320 μm. Figure  4 shows SPR images of mouse broblast cells (3T3) seeded on gold coverslips, showing the wide FOV and the potential to image relatively large populations of cells, for example, there are ~40 cells within gure 4. 3.1.2. Spatial resolution. The resolution of a conventional light microscope is limited by the diffraction limit [ 42 ]. The theoretical diffraction-limited resolution of the optical imaging modalities presented (transmission and re ection microscopy) for a 60×, 1.49 NA microscope objective with 680 nm incident light given by the Abbe limit is 0.28 μm. However, by under magnifying to 18×, the minimum resolvable feature, magni ed onto a 2 × 2 region of the sensor in accordance with the Nyquist criteria, is 1.1 μm. The average measured size of a typical cultured mammalian neuronal soma is ~16 μm (data not shown) [ 43 ] and a typical cardiomyocyte cell is about 100 μm long and 10–25 μm in diameter [ 44 ]. Even after undersampling the resolution is still more than adequate for imaging both cardiomyocytes and networks of neurons including many of the larger dendritic and axonal processes which are typically 1–2 μm in size [ 45 ].

The resolution in SPR sensors, however, is determined by the propagation length of the SPs. The plasmons propagate along the metal/dielectric interface for a distance decided by the propagation length before they decay back into photons.

This limits the resolution of the system in the direction parallel to the propagation of SP waves. So, when the propagating plasmon reaches a refractive index boundary it is no longer supported and is re-radiated as a photon. This means that even though in theory the plasmon propagation length in water on a 50 nm gold lm is typically ~3.1 μm [ 18 ], ne details, such as neuronal axons and dendrites can be resolved. Figure 5(c) shows a primary, hippocampal neuron cultured on a gold sensor imaged at 5 d in vitro (DIV). The ne, elaborating features visible in the gure demonstrate that axons and dendrites can be clearly resolved using SPR imaging in this regime. However, cell processes that are perpendicular to the propagation of the SPR are hardly resolved, due to the resolution limit of the SPR system, while they are clearly visible when they are parallel to the propagation of the SP wave [ 26 ].

To demonstrate the resolution limitations of the SPR microscopy, we explored an alternative evanescent wave microscopy technique that is based on total internal re ection ( gure 5(b)). Primary hippocampal neurons were cultured on glass coverslips treated with PLL. The glass surface was illuminated with a uniform disc of light from the 640 nm LED source objective BFP. This disc of light was stopped down using the aperture diaphragm (D1) to just above the total internal re ection critical angle to gain contrast.

Figure 5(b) shows that TIRM can resolve dendrites in both directions demonstrating a superior resolution compared to SPR. Although TIRM provides resolution and contrast based on the frustrated total internal re ection, it has a limited ability to resolve functional time-resolved information compared to SPR. 3.1.3. Refractive index mapping. In biological samples, spatial differences in the lipids and proteins present result in a variation of the refractive index. These differences provide the contrast in an SPR image because as previously explained, refractive index discontinuities in the sample result in the propagating plasmons being re-radiated as photons.

To assess the spatial variation of the refractive index on our gold surfaces, the refractive index of each pixel in the FOV of gure  5(d) was determined by scanning the illumination angle, gure 6(a) and nding the angle of the minimum re ection for each pixel, gures 6(b) and (c). Then by solving the Fresnel equations  for re ection and transmission in a multilayer device, the refractive index was calculated (see supplementary information).

Figure 6(a) shows the full image at various angles during the scan. The re ection as a function of scan angle at one x-position and one y-position was determined and shown in gure 6(b). The dark line in the projections represents the SPR dip and corresponds to the angle of incidence at the gold layer. The complete angle scan is shown at various points on the sample in gure  6(c). Note that at different angles of incidence plasmons are excited successively either under the cells or on cell-free background areas.

Figure 6(d) shows that the effective refractive index is reasonably homogeneous underneath the soma indicating the cell has adhered to the surface uniformly. The effective refractive index in a multilayer system is a function of the gap distance between the SPR sensor and the cell membrane and is weighted with an exponential decay [ 8, 10 ]. A larger effective refractive index indicates that the cell membrane has closely adhered to the surface, conversely a smaller measured refractive index indicates that the cell membrane is further away from the surface. Cell membranes are composed mostly of lipids and proteins which have refractive indices in the range of 1.46–1.54, so the refractive index we have measured here of 1.40 for the cell structure is less than previously reported values [ 46, 47 ]. The lower RI can be attributed to the lower RI cytosol (n = 1.35 [ 46 ]) reducing the effective RI alongside the gap distance. It is important that the gap distance is less than the evanescent eld distance and this will be addressed next. 3.2. SPR-excited uorescence imaging As described in the introduction, the evanescent eld generated at the interface can excite uorophores placed within it. Exploiting this phenomenon, we were able to determine whether the neurons were close enough to the gold surface for their membranes to lie within the evanescent eld by patching a microelectrode lled with dye into the cell. The dye within the cell was excited using both epi- and SPR-excited uorescence, gures 7(a) and (b), respectively. Figure 7(b) con rms that the 680 nm SLED SPR system can excite the dye-labeled cell and therefore, the cell is within the evanescent eld. This is a necessary prerequisite to allow functional imaging using

SPR. 100% of cells tested (N = 12) were found to lie within the evanescent eld. This demonstrates that SPR can be used to excite either structural or functional uorescent dyes in studies of neuronal (or other) cultures.

For label-free functional imaging of dynamic, cellular functions from living cells, a low noise, high sensitvity system is required. We characterized this sensitivity by changing the bulk refractive index of the dielectric on a gold surface. This characterization is followed by results that demonstrate the ability of the system to image localized dynamic processes by detecting spatial variations in the contraction of stem cellderived cardiomyocytes. 4.1. Sensitivity to refractive index changes The sensitivity of the gold surfaces on the SPM was experimentally determined by altering the bulk refractive index of the dielectric on a 50 nm planar gold sample.

The concentration of the solution on the surface on the substrate was changed to either increase or decrease the refractive index using NaCl solutions. The concentration values were chosen so that the refractive index increase or decrease was in step sizes of Δn = 0.1 × 10–3. To nd the sensitivity, the angle of incidence was chosen such that the re ection gradient (ΔR) is at its maximum. Practically, it is dif cult to set this parameter, so the dynamic range was rst established by angle-scanning across the SPR dip, and setting the SPR illumination angle to the location of the greatest gradient ( gure 3(b)). There is a trade-off between greater sensitivity and SNR, as at the angle of the maximum gradient the light intensity is much lower, reducing the Poisson-limited SNR. Considering this, the angle of incidence was xed at around 30% of the dynamic range.

The results from all the experiments are summarized in gure  8. Figure  8 shows that the normalized light intensity (ΔR) increases linearly with increasing refractive index. The sensitivity of the sensor was characterized by the RIU as 2 × 10−5.

The experimentally measured sensitivity here is less than the value of ~10−6 RIU reported previously from intensity modulating SPR techniques in literature [ 21, 22 ]. However, the level of detection in these sensitivity experiments is limited by the exposure time because the power of the noise is proportional to the bandwidth. So, for the relatively small bandwidth used for these measurements (286 Hz), the total noise power may be reduced if a lower sample frequency is used thus increasing the SNR. 4.2. Detection of cardiomyocyte contraction In order to demonstrate a potential experimental scenario where a multimodal system would be bene cial to the user, hESC-cardiomyocytes were cultured on gold surfaces. First, bright- eld microscopy was used to locate a hESC-CM; see gure 9(a). The periodic contractions of a living, beating cell could be visualized by monitoring the change in SPR intensity on the EPIX camera. By taking the difference of each subsequent frame with the rst frame, a stack was produced that shows the area where there is the greatest variation in light intensity (‘hot-spot’); see gures 9(c) and (d). The background was corrected for by taking the difference of the ‘hot-spot’ with a sample of the rest of the cell which was not moving.

Figure 9(e) shows that the SPR intensity changes in time with the contractions of the stem cell-derived cardiomyocyte (‘signal’) with a high signal-to-noise ratio. Two control ROIs are plotted. One on the background and a secondary location within the cell. The resulting SPR responses show that only a couple of areas are moving within the cell, with the rest being uniform. The area where there is the greatest change in light intensity appears to be around the nucleus of the cell (see supplementary video 1). It is possible that the source of the signal is from the stress bres or contractile apparatus within the hESCCM pressing on the nucleus causing it to move. The observed contraction localized in the centre of the cell is typical for hESC-CM’s attached to rigid surfaces [ 48 ]. The control ROI on the ‘background’ was taken directly next to the beating cardiomyocyte, which shows no change in relative SPR intensity. This demonstrates that the change in SPR intensity is localized to the cell of interest and therefore, most likely due to refractive index changes or localized movement within the cell and not from movement causing mechanical waves throughout the volume. To fully utilize the multi-modal system, uorescent labelling could be used for comparison of the label-free results with well-established labelling techniques.

In this paper, we have presented a multimodal platform for cell physiology combining SPR imaging with a number of ancillary microscopy systems. We have shown that the system is capable of both structural and functional imaging of cultured cells. Using the system for structural imaging, a number of modalities can be exploited including re ection/ transmission microscopy, TIR micrscopy, epi- uorescence and SPR imaging to obtain complementary information. A wide eld-of-view has been demonstrated with a suitable spatial resolution for imaging cardiomyocytes and resolving individual axons and dendrites in cultured primary neurons label-free.

The ability of the system to study spatiotemporal cellular functions was demonstrated by imaging localized contractions of stem cell-derived cardiomyocytes. To the best of the author’s knowledge, this is the rst demonstration of functional imaging of the refractive index changes using SPR in single cardiomyocytes. Using SPR could allow the localized contractions of cardiomyocytes to be imaged in real-time and drugs to be tested in vitro providing additional information compared to traditional non-imaging techniques [ 49, 50 ] and less computation compared to video microscopy [ 51 ].

In future work, a small network of cultured neurons will be grown on the surface of a sensor and any small changes to this light that occur during an action potential will be monitored.

The authors would like to thank Tim Smith, Jas Chauhan, and Dave Taylor for their technical assistance. This work was supported by the Engineering and Physical Sciences Research Council (Grant Nos. EP/F005512/1, EP/H022112/1, EP/ G058121/1), the last grant being a Royal Academy of Engineering/EPSRC Post-doctoral Fellowship to K F W; the British Heart Foundation (Grant Nos. SP/15/9/31605, RG/15/6/31436, PG/14/59/31000, RG/14/1/30588, RM/13/30157, P47352/ CRM); BIRAX (Grant No. 04BX14CDLG); the Medical Research Council (MRC) (Grant No. MR/M017354/1); the National Centre for the Replacement, Re nement & Reduction of Animals in Research (NC3Rs) (Grant Nos. CRACKIT:35911-259146, NC/K000225/1); and The Whitworth Society, providing nancial support to C L H.