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Abstract
We present a new approach for three-dimensional (3D) live single-cell imaging with isotropic sub-micron spatial resolution using fluorescence computed tomography (fCT). A thin, highly inclined and laminated optical (HILO) sheet of light is used for fluorescence excitation in live single cells that are rotated around an axis perpendicular to the optical axis. During a full rotation, 400-500 two-dimensional (2D) projection images of the cell are acquired from multiple viewing perspectives by rapidly scanning the HILO light sheet along the optical axis. We report technical characteristics of the HILO approach and the results of a quantitative comparison with conventional epi fCT, demonstrating that HILO fCT offers significantly (about 17 times) reduced photobleaching and a two-fold improvement in 3D imaging contrast. We discuss potential application areas of the method for cell structure studies in live single cells with isotropic 3D spatial resolution.
© 2018 Optical Society of America under the terms of the OSA Open Access Publishing Agreement
1. Introduction
Three-dimensional (3D) organization of a cell is closely associated with its function [1, 2]. Quantitative imaging of cell behavior and its architectural dynamics in 3D are therefore essential for understanding cellular function in normal and disease states [3–5]. Absolute measurements of the cellular structure in 3D require the spatial resolution of images to be the same in all three Cartesian directions (isotropic). However, the vast majority of 3D imaging techniques are based on optical sectioning, a method that is limited by anisotropic spatial resolution, roughly two times lower in the axial (along the optical (Z) axis) as compared to the lateral (in the imaging (XY) plane) direction. To overcome this limitation, several advanced optical imaging methods based on optical sectioning have been developed. For instance, the I5M technique uses two opposing objective lenses to observe and illuminate from opposite sides of the sample simultaneously to improve the axial resolution to 70 nm [6]. However, due to the requirement of maintaining the two objective lenses aligned in 3D space with sub-micron accuracy during image acquisition, the utilization of this technique has been limited to specialized research labs [7]. Dual inverted selective-plane illumination microscopy (diSPIM) [8], is another method that approximates isotropic 3D spatial resolution and is implemented by alternating the illumination and detection optical paths between two orthogonally oriented objective lenses. However, because of the geometrical constraints of this configuration, the method is limited to the use of excitation objective lenses with relatively low numerical apertures and inferior spatial resolution. Moreover, the axial resolution of the above methods is different from the lateral resolution; therefore, both methods cannot provide true 3D isotropic optical imaging.
An alternative optical imaging technique that offers truly isotropic diffraction-limited 3D spatial resolution is based on the principles of computed tomography (CT) [9–19], as opposed to optical sectioning. Similar to X-ray CT, fluorescence computed tomography (fCT) utilizes a series of 2D projection images of the cell taken from multiple perspectives to computationally reconstruct a volumetric 3D image with isotropic spatial resolution. Typically, 300-500 projection images from different orientations of the specimen are acquired to ensure adequate spatial sampling of the object. Several research groups, including ours, have exploited optical CT for 3D imaging of hematoxylin and eosin stained fixed cells in absorption mode [12, 18]. In a previous study, we reported a conventional epi fCT method for imaging live cells (which is called epi Live-Cell CT or eLCCT) with isotropic, diffraction-limited 3D spatial resolution [17]. Because eLCCT is based on the use of epi fluorescence illumination, it suffers from reduced image contrast due to the partial inclusion of out-of-focus light that is collected during imaging. Another limitation is photobleaching and potential phototoxicity when imaging live cells—a result of continuous excitation of fluorescence dyes across the entire cell during the image acquisition process. We recognized that photobleaching is a limiting factor of the eLCCT technique, especially when performing time-lapse imaging.
Light-sheet microcopy is a cutting-edge optical imaging method to decrease photobleaching and phototoxicity [20, 21]. However, traditional light-sheet microscopy utilizes two orthogonal objective lenses for excitation and detection, independently, and is incompatible with our eLCCT system. Jurgen Mayer et al. combined the traditional light sheet microscopy (SPIM) with optical projection tomography, but the resolution is limited to 3-4 μm [22]. Highly inclined and laminated optical (HILO) microscopy [23], which utilizes only one objective lens to generate a micron-thin optical sheet for excitation and utilizes the same objective lens for detection, can be integrated with our eLCCT. Therefore, to address our eLCCT limitations, we developed a new fCT method based on HILO combined with LCCT (which is called HILO LCCT or hLCCT) for quantitative 3D imaging. To this end, we employed a laser beam as a fluorescence excitation source that is focused on the back focal plane (BFP) of the imaging objective lens and offset from the optical axis to produce HILO illumination. In contrast to epi illumination, where the entire cell is illuminated during image acquisition, HILO illumination is restricted to a thin (1-2 µm) section of the cell at a time, thus limiting the excitation of the fluorophores to a volume near the focal plane. For multi-perspective imaging, the cell is rotated continuously at a rate of 2-3 rpm by means of high frequency electrical fields in a microfluidic chip with a rotation axis perpendicular to the optical axis of the imaging system [17]. To obtain a 2D projection image, the illuminating optical sheet is rapidly scanned along the optical axis across the entire cell while keeping the camera shutter open. Using appropriate computational algorithms, a volumetric 3D image of the cell is then reconstructed from a series of 2D projection images taken over a full rotation of the cell. We demonstrate that our approach offers a two-fold higher imaging contrast and approximately 17-fold reduced photobleaching as compared with conventional epi fCT.
2. Results
2.1 System design
The hLCCT imaging system (Fig. 1(a)), is built around an inverted epi fluorescence microscope and uses a fiber-coupled laser as the illumination source that is coupled into the microscope through a modified total internal reflection fluorescence (TIRF) module, which is mounted on an epi illumination mount. The TIRF module comprises two achromatic lenses (L1, L2) to focus the laser beam on the BFP of the objective lens. A multichroic mirror is mounted in a filter cube and placed inside of the corresponding microscope filter turret to deflect the laser beam onto the objective lens for fluorescence excitation. The lateral position of the focal point of the laser beam at the BFP of the objective lens (OL) can be adjusted with respect to the optical axis by two alignment knobs on the TIRF module. The laser beam is offset 2.6 mm away from the center of the optical axis. As a result of passing through the objective lens, the laser beam becomes effectively an inclined sheet of light with respect to the imaging plane (Fig. 1(b)). The objective lens, which is mounted on a scanning piezoelectric stage, is rapidly swept along the optical axis, with the focal plane and illumination sheet traversing the entire cell in tandem to obtain one 2D projection image of the cell per sweep. Consequently, a series of projection images from multiple perspectives of the cell over one full rotation are acquired. Although several other cell rotation methods have been exploited [24, 25], a customized microfabricated “electrocage” chip is used for cell rotation, as described elsewhere [17, 26]. Briefly, cells are delivered to the rotation area through a microfluidic channel on the electrocage chip. A pipette tip is connected to the inlet and acts as a reservoir of cells, and a 10 μL syringe is connected to the outlet to aspirate the cells into the cage. A triple-band bandpass emission filter is placed in the emission path to filter out the residual excitation light. A cooled electron-multiplying charge-coupled device (EMCCD) camera is connected to the right port of the microscope for acquiring 2D projection images. The piezo stage, EMCCD, and the laser are synchronized via custom software developed in-house. A swept field confocal imaging module is mounted on the left port of the microscope for comparison purposes with hLCCT. In addition, to compare the hLCCT with our previously developed eLCCT under the same imaging conditions, a multispectral LED light source (not shown in Fig. 1(a)) is mounted on the second Nikon epi-fl illumination mount of the same microscope as the light source for eLCCT. By placing the filter cube into the corresponding filter turret, we can switch between the eLCCT and hLCCT easily. Both the eLCCT and hLCCT system share the same optical detection path, and the optical pixel size of the image taken by eLCCT and hLCCT is 107 x 107 nm. Detailed information about the system components is provided in the ‘Materials and Methods’ section. Figure 1(b) shows a detailed schematic of the HILO illumination on a single cell.
Fig. 1 (a) Optical system setup of hLCCT. FL: fiber-coupled laser; L1, L2: achromatic lenses; MM: multichroic mirror; EF: emission filter; M: mirror; TL: tube lens; EMCCD: cooled electron-multiplying charge-coupled device; PZT: piezoelectric transducer; OL: objective lens; MS: microscope stage; ECC: electrocage chip; PT: pipette tip; S: syringe; RA: rotation area (electrocage). The illumination beam and the fluorescence are shown in purple and blue, respectively. (b) A detailed schematic of the HILO illumination. R is the diameter of the field of view (FOV). The calculated incident angle is 77.8°, and the thickness of the HILO illumination is 1.7 μm.
2.2 HILO optical sheet thickness
We first investigated the uniformity of the HILO beam in 3D space over the extent of the cell and assessed its influence on the projection image contrast and the photobleaching rate of a fluorescent dye. Furthermore, to acquire 2D projection images of the cell, the focal plane of the objective lens needs to be scanned across the entire cell along the Z-axis. Owing to the fixed location of the laser beam focus in the optical path, scanning the objective lens may cause the illumination sheet to converge or diverge at the focal plane, resulting in optical sheet thickness variations across the cell, which would increase photobleaching and decrease the imaging contrast.
According to [23], the thickness of the HILO beam, dz, can be calculated by using the following equation:
where R is the diameter of the field of view (FOV), and θ is the incidence angle of the optical sheet. The incidence angle θ can be calculated fromwhere x = 2.6 mm is the distance of the focal point of the excitation laser beam from the optical axis at the BFP, n = 1.33 is the refractive index of the cell medium, and fobj = 2 mm (100x objective lens) is the focal length of the objective lens. In our case, the diameter of the FOV is 8 μm when the objective lens is located at Z = 0, resulting in a calculated incidence angle θ and HILO beam thickness of 77.8° and 1.7 μm, respectively.In this work, we measured the uniformity of the thickness of the HILO optical sheet using 200 nm fluorescent polystyrene beads and recorded the emission intensity of the beads as a function of the location of the objective lens along the optical Z-axis at different XY positions within the FOV. In each measurement, we first fixed the objective lens at the desired location Zobj, then a cover slip with 200 nm diameter fluorescent beads was placed on a motorized microscope Z stage (NanoScanZ, Prior Scientific, Rockland, MA), and moved across the HILO beam with a step size of 100 nm in the Z direction. Images of the beads were recorded by the camera, and the optical sheet thickness was determined as the full-width at half-maximum (FWHM) of the bead intensity profile. At each objective lens location Zobj, we measured the thickness of the HILO beam at the center (X = 0, Y = 0) and the other four locations near the corners of the FOV (point 1: X = 4 μm, Y = 4 μm; point 2: X = 4 μm, Y = −4 μm; point 3: X = −4 μm, Y = 4 μm; point 4: X = −4 μm, Y = −4 μm). Second, while at each XY location, we translated the objective lens over 10 μm below and above the starting position (Zobj = 0 µm) along the Z axis with a step size of 2 μm. The measurements were repeated with five different single beads to assess experimental error and provide a more accurate measurement result in the center of the FOV, and one single bead was used to characterize the HILO beam thickness at the four corner points. The detailed experimental setup is shown in Fig. 2(b), and the measurement results are shown in Fig. 2(c) and Fig. 2(d). A representative measurement demonstrating the thickness of the HILO sheet at the center of the imaging plane (X = 0, Y = 0) when the objective lens was located at Zobj = 70 µm is shown in Fig. 2(e).
Fig. 2 (a) Changes in HILO beam thickness at different locations of the objective lens along the optical axis; (b) Experimental setup for measuring the HILO beam thickness; Quantitative measurement of the HILO beam thickness at the center (c) and four corners (d) of the FOV, as a function of the objective lens position along the optical Z axis. Five independent measurements were performed to characterize the thickness in the center of the FOV, whereas one measurement was used for the beam thickness in each of the four corners. The error bars in (c) represent the standard deviation of the thickness measurements obtained with five independent measurements with different beads at each Zobj location, the variation of the error bars at different objective lens Zobj position may be caused by the PZT electrical noise. The black dot indicates the average thickness of the HILO beam at the corresponding Zobj location. A thickness change in (d) at Zobj = 2 μm at the corner point X = −4 μm, Y = 4 μm is 0.25 μm, and is within the measurement error. (e) Representative intensity profiles of the beads at the center of the imaging plane when the objective lens was located at 70 µm. The result was averaged over five different single beads. Zb indicates the beads location along the optical axis.
We observed that the average thickness of the HILO beam at the center of the FOV as a function of Z position ranged from 1.66 to 1.77 μm. The average thickness at the center of the FOV across the entire scan range of 20 μm was 1.69 ± 0.17 μm, which is close to the theoretical value of 1.67 μm. The beam thickness at the four corners varied from 1.5 µm to 1.9 µm. Compared with the thickness in the center, the maximum difference in the corners was 0.21 µm or 12% of the average thickness at the center of the FOV. We note that the sheet thickness increases by 0.25 μm from Zobj = 0 μm to Zobj = 2 μm, which represents ~15% relative change and lies within measurement errors.
2.3 Photobleaching
To quantitate photobleaching, we performed a series of experiments where we measured the fluorescence emission intensity as a function of illumination time. Mitochondria of two sets of Barrett’s esophagus metaplastic epithelial cells (CP-A cell line), each containing 6 cells, were fluorescently stained and excited continuously for 300s at an imaging rate of 10Hz, using eLCCT or hLCCT. The objective lens was continuously scanned across the cell from −10 µm to + 10 µm to acquire a number of 2D projection images over time. To compare the photobeaching rate under the same conditions, the power of the fiber laser used in the hLCCT method and the LED source used in the eLCCT method were adjusted to ensure that the average intensity of the region of interest in the acquired projection images was approximately the same in both cases. The same multichroic mirror was used in both the hLCCT and eLCCT systems to ensure the excitation and emission wavelength are the same. The two sets of six CP-A cells were imaged and the photobleaching rate was calculated and fit by using an exponential decay function y = y0 + Aexp (-R0t), where y0 is a constant, t is the time and R0 is the decay rate. The emission intensity levels in the hLCCT and eLCCT modes were measured and characterized as a function of time, and the result is shown in Fig. 3(a). As expected, we observed a markedly decreased photobleaching rate in hLCCT (4.2 × 10−4) as compared with eLCCT (7.2 × 10−3), i.e. 17.1 times less than that in eLCCT. A representative comparison example of photobleaching by using hLCCT and eLCCT is shown in Fig. 3(b) and Fig. 3(c), respectively.
Fig. 3 (a) Comparison of photobleaching kinetics between hLCCT (black squares) and eLCCT (red circles). The measurements were fit with an exponential decay function (solid curves). The fluorescence intensity decay rate observed with hLCCT was 17.1 times lower than eLCCT. (b) and (c) A comparison of photobleaching observed with hLCCT and eLCCT. Mitochondria of two different CP-A cells were fluorescently stained with Mitotracker Red. The cells were continuously illuminated for 300 s in both cases. The imaging conditions were set to provide comparable maximum fluorescence signals by adjusting the excitation power. The fluorescence emission intensity was calculated by averaging the signal of a feature inside of the solid yellow and red box in hLCCT and eLCCT, respectively.
2.4 Contrast
To characterize the contrast of hLCCT and eLCCT, we compared reconstructed 3D images of the same cell nucleus obtained using both methods. The characterization of contrast was performed using fluorescently labeled nuclei of five different K562 cells. The cells were rotated in the electrocage, and their 2D projection images were acquired using the hLCCT and eLCCT methods, then the 3D images of the cells were reconstructed computationally. We used the Michelson contrast (visibility), which is defined as (Imax-Imin)/(Imax + Imin), as a measure to quantify the contrast. To this end, we measured the average intensity of an arbitrarily chosen bright feature (Imax) and a dark area with low signal (Imin) within the cell nucleus in the reconstructed 3D images generated using both methods. The image contrast of five K562 cell nuclei were determined to be 33%, 35%, 22%, 22% and 45% with hLCCT, compared with 16%, 23%, 12%, 12% and 21% with eLCCT, respectively, an approximately two times increase in hLCCT. The detailed comparison result of five cells is shown in Fig. 4(a). Figure 4(b) demonstrates one representative contrast comparison using 3D images of the same K562 cell. Imax is the average intensity in the red box, and the Imin is the average intensity in the blue box. The contrast of eLCCT and hLCCT of this measurement is 20.63% and 44.64%, respectively, demonstrating the 3D image contrast of hLCCT is about two-fold higher than that using eLCCT.
Fig. 4 (a) Contrast comparison of reconstructed 3D images of five K562 cell nuclei obtained using the hLCCT and eLCCT imaging methods. The contrast is two or more times higher in hLCCT than that obtained with eLCCT. (b) A representative example of 3D cell images contrast comparison using the eLCCT and hLCCT methods. The contrast is computed using the average intensities in the red box as Imax and blue box as Imin. The contrast of the 3D cell image acquired with hLCCT is enhanced two-fold compared with eLCCT.
2.5 Comparison with confocal and eLCCT 3D imaging of live single cells
For comparison, we acquired 3D images of the same nucleus using hLCCT, eLCCT, and confocal imaging. To acquire a z-stack with confocal imaging, the cell was first held steady in the electrocage. Afterwards, the cell was rotated for the two other imaging modalities. Our results, which are shown in Fig. 5(a) and Visualization 1 for eLCCT, Fig. 5(b) and Visualization 2 for hLCCT, demonstrate isotropic 3D spatial resolution obtained with both LCCT modalities, as evidenced by the absence of any image distortion when observing the cell from two orthogonal orientations. In contrast, the confocal image, which is shown in Fig. 5(c) and Visualization 3, shows marked blurring and elongation along the optical axis owing to the inferior resolution along the z (optical) axis inherent to conventional confocal microscopy.
Fig. 5 Comparison of (a) eLCCT, (b) hLCCT, and (c) confocal images of a K562 cell nucleus. Maximum intensity projections of a 3D reconstructed volumetric image showing the nucleus in two orthogonal orientations (0° - upper panel, 90° - lower panel) generated with the corresponding imaging modalities. Scale bars: 5 µm. Visualization 1, Visualization 2, and Visualization 3 demonstrate the reconstructed 3D images of the nucleus in case of eLCCT, hLCCT, and confocal Z-stacks.
3. Discussion and conclusion
In this paper, we present a new fluorescence computed tomography approach based on HILO illumination applied to 3D imaging of single live cells with isotropic spatial resolution. The approach (hLCCT) represents an improvement in photobleaching and image contrast over the previously reported eLCCT method [17], whereby the HILO optical sheet replaces the conventional epi fluorescence excitation. Our analysis revealed that the HILO beam is thin and reasonably uniform across the field of view. Furthermore, the beam mainly retains its intensity profile when the objective lens is scanned along the Z-axis over distances equivalent to cell dimensions. Compared with eLCCT, the hLCCT approach enhances the 3D image contrast about two times, while decreasing the photobleaching rate by 17-fold. In combination with the diffraction-limited 3D spatial resolution offered by the fCT method, our approach facilitates long-term fluorescence imaging in biomedical research, such as monitoring protein interaction with cellular compartment-dynamics in living cells [27, 28] or cellular architecture dynamics under physiological conditions [29].
The NA of the objective lens primarily determines the spatial resolution of hLCCT, which is the same as the eLCCT and diffraction-limited. Other factors, such as the number of projection images acquired per full rotation, cell rotation stability etc., can also affect the resolution of hLCCT. In our eLCCT paper, we have investigated that the average lateral shifts of a 10 μm diameter cell in the x and y directions during a full rotation are 0.057 μm and 0.84 μm, which will decrease the resolution 0.76 nm and 1.12 nm respectively. Furthermore, we have measured that the averaged cell rotation instability is about 5.1% per full rotation; this will cause a 42 nm resolution decrease of a 10-μm dimeter cell. Compared with the eLCCT system resolution (280 nm to 290 nm at 550 nm excitation laser), the resolution decreased by the cell rotation instability is small and can be neglected. The current temporal resolution of our hLCCT system is 20-30 s and is limited by the scanning speed of the objective lens. The imaging speed can be improved using extended depth-of-field techniques such as wavefront coding [30–32], though at the cost of increased complexity of image reconstruction. Another drawback of our hLCCT system resides in the relatively narrow field of view (8 μm in diameter) compared with eLCCT (54 μm x 54 μm) due to inclined beam illumination on the sample, which lies at an angle to the imaging plane, resulting in blurring regions of the sample away from the designed focal plane. However, this problem can be solved by the use of oblique plane microscopy [33], which refocuses the blurred regions of the sample by tilting the imaging plane, at the cost of increased complexity of the optical system.
In summary, we demonstrated a new fCT approach for 3D imaging of live single cells based on HILO illumination. The approach has diffraction-limited true isotropic 3D spatial resolution, improved image contrast, and more than an order of magnitude decreased photobleaching. These features make the presented method a powerful tool in the biomedical imaging field.
4. Materials and methods
4.1 Hardware specifications
The components used in the hLCCT setup are as follows: a Nikon microscope (Eclipse Ti-U, Nikon, Melville, NY); a Nikon TIRF module (T-FL-TIRF2, Nikon); a motorized microscope stage (MS-2000, Applied Scientific Instrumentation, Eugene, OR); a fiber-coupled laser (Aurora, Bruker, Middleton, WI); two achromatic lens in Nikon TIRF2 (L1: 30 mm Dia. x 10 mm focal length, VIS-NIR Coated, Achromatic Lens; L2: 30 mm Dia. x 250 mm focal length, VIS-NIR Coated, Achromatic Lens, Thorlabs, Newton, NJ, USA); a triple-band mirror (FF436/514/604-Di01-25x36, Semrock, Rochester, NY); an EMCCD (Evolve 512, Photometrics, Tucson, AZ); a pipette tip (10 μL, Fisherbrand SureOne Micropoint Pipette Tips, ThermoFisher Scientific, Waltham, MA); a PZT (SFC-CTL V1, Bruker); an objective lens (S-Fluor 100x oil immersion objective lens, NA 1.3, Nikon); a 10 μL syringe (Kloehn pump, Las Vegas, NV); a swept field confocal (SFC, Bruker); a multispectral LED light source (SpectraX 6-NII-SE, Lumencor, Beaverton, OR); a triple-band emission filter (FF01-457/530/628-25, Semrock).
4.2 Cell culture
Immortalized human chronic myelogenous leukemia (K562 cell line) cells were cultured in RPMI 1640 medium (Gibco, Grand Island, NY, USA) containing 10% fetal bovine serum, 100 units/mL penicillin, and 100 µg/mL streptomycin at 37°C in a humidified incubator containing 5% CO2. The cell density was determined using a Countess II FL Automated Cell Counter (Life Technologies, Carlsbad, CA). Cultures were maintained by the addition or replacement of fresh medium.
CP-A cells were cultured using Gibco keratinocyte serum-free cell growth medium (10724-011, Invitrogen, Carlsbad, CA), supplemented with human EGF at 2.5 µg/500 mL, bovine pituitary extract (BPE) at 25 mg/500 mL and penicillin–streptomycin solution at 100 µg/mL (15140122, ThermoFisher Scientific). CP-A cells were cultured at 37 °C in a humidified incubator containing 5% CO2 and passaged upon reaching 70 - 80% confluency, once to twice weekly. To minimize over-trypsinization, passaging was performed using 0.05% Trypsin-EDTA (25300-062, ThermoFisher Scientific).
4.3. Fluorescent beads sample preparation
The 200 nm beads were prepared by diluting 1 μL of bead stock solution (F-8767, 505/515, ThermoFisher Scientific) in 10 mL 90% methanol to decrease beads surface tension and prevent aggregation. The beads solution was vortexed for 1 min and checked under a fluorescence microscope afterwards to ensure there were no large bead aggregates present and that single beads were sparsely spread in the field of view. To prepare a sample for measuring the thickness of the HILO laser beam, 5 μL of the diluted beads solution was transferred onto a clean microscope slide, spread with a pipette tip and allowed to air-dry. The dried beads were covered with a cover slip and sealed around the edges with glue.
4.4 Cell nucleus, mitochondria and beads imaging
The nuclei of the K562 cells were stained with Hoechst 33342 (H3570, ThermoFisher Scientific), and excited with fluorescence at 405 nm (hLCCT) or 395 nm (eLCCT). The mitochondria of CP-A cells were stained with Mitotracker Red (1M7512, ThermoFisher Scientific). The excitation wavelength used for mitochondria fluorescence excitation was 561 nm (hLCCT) or 555 nm (eLCCT). The excitation wavelength used for the fluorescent beads was 488 nm. The exposure time was 40 ms and the PZT scan period was 100 ms. No excitation filters were used for all the cell nuclear, mitochondrial, and beads imaging. The triple-edge dichroic mirror was used to separate the excitation and emission beam, and the triple-band bandpass emission filter was used to block the residual excitation light. In confocal imaging, the light source was a fiber-coupled laser, which was the same as used in hLCCT, the excitation wavelength for the K562 nuclei was 405 nm. A notch filter (488/561/640) was used as the emission filter, the step size in Z was 0.5 μm, and the exposure time was 40 ms.
4.5 Cell rotation
A customized microfabricated “electrocage” chip, which has eight microelectrodes arranged on two layers and a microfluidic channel, was used to rotate cells. A function generator (PXI-1045, National Instruments Corporation, Austin, TX) was used to create sinusoidal waveforms to drive the electrocage; a customized interactive LabVIEW program was developed to control the eight individual channels of the PXI, enabling the user to modify the peak-to-peak voltage, frequency and phase of the PXI channels. The electrocage chip fabrication, cell rotation and characterization used in hLCCT were identical to the eLCCT system and can be found in [17].
4.6 3D imaging reconstruction
The procedure of 3D imaging reconstruction in hLCCT method is the same as in eLCCT method [17].
4.7 Determination of Field of View (FOV) in hLCCT method
The FOV of our hLCCT system was determined as the distance between the locations where the bead intensity drops to half of its maximum intensity on two opposite sides of the imaging plane by moving a 200 nm fluorescence bead with respect to the imaging plane (Fig. 6). The diameter of the FOV can be recorded and obtained by the digital controller of the motorized stage. The diameter of the FOV is 8 μm when the objective lens was located at Z = 0 µm (starting position).
Funding
This work was supported by a grant from the W. M. Keck Foundation (024333-001) to D.R.M.
Acknowledgements
The authors thank Sandhya Gangaraju and Morgan Bennett for cell culture needs. This work was supported by a grant from the W. M. Keck Foundation to D.R.M.
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