Reflection phase microscopy by successive accumulation of

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Reflection phase microscopy by successive accumulation of interferograms Min Gyu Hyeon, Taeseok Daniel Yang, Jin-Sung Park, Kwanjun Park, Yong Guk Kang, BeopMin Kim, and Youngwoon Choi ACS Photonics, Just Accepted Manuscript • DOI: 10.1021/acsphotonics.8b01703 • Publication Date (Web): 12 Feb 2019 Downloaded from http://pubs.acs.org on February 13, 2019

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Reflection phase microscopy by successive accumulation of interferograms Min Gyu Hyeon1, Taeseok Daniel Yang2, Jin-Sung Park3, Kwanjun Park4, Yong Guk Kang4, Beop-Min Kim1,2,4*, and Youngwoon Choi2,4* 1Department

of Biomicro System Technology, Korea University, Seoul 02841, Korea of Biomedical Engineering, Korea University, Seoul 02841, Korea 3Center for Molecular Spectroscopy and Dynamics, Institute for Basic Science (IBS), Seoul 02841, Korea 4Department of Bio-convergence Engineering, Korea University, Seoul 02841, Korea *Correspondence: [email protected], [email protected] 2School

Abstract Imaging 3-dimensional (3-D) structures of biological specimens without exogenous contrast agents is desired in biological and medical science in order not to disturb the physiological status of the living samples. Reflection phase microscopy based on the interferometric detection has been useful for the label-free observation of such samples. However, the achievement of optical sectioning has been mainly based on the time gating set by the broad spectra of light sources. Here we propose wide-field reflection phase microscopy using a light source of narrow bandwidth, which is yet capable of achieving the optical sectioning sufficient for 3-D imaging of biological specimens. The depth-selectivity is achieved by successive accumulation of interferograms (SAI) produced by synchronous angular scanning of a plane wave on both the sample and reference planes. This intensity-based cumulative process eventually results in a coherent addition of object fields that quickly attenuates the out-of-focus information along the axial direction. We theoretically investigated and numerically verified the generation of the depth-selectivity by SAI. We also implemented a reflection phase microscope working with this principle, and then demonstrated high-resolution 3-D imaging of living cells and small worms in the label-free manner.

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KEYWORDS: quantitative phase imaging, synchronous angular scanning, depth selectivity, 3-D imaging, live cell imaging, cellular membrane fluctuation

Observation of objects in 3-D space with high spatial resolution has been desired in wideranging areas, such as biology, medicine, physical science, and industrial applications. Confocal microscopy 1-3 and multi-photon microscopy 4-7 have proven to be useful for highresolution 3-D imaging, especially for biological specimens. Usually, 2-dimensional (2-D) images are acquired by scanning a tightly focused laser beam, and a 3-D map is produced by combining the multiple 2-D images taken at different target depths. These imaging methodologies have shown high resolution along the longitudinal directions, as well as superior optical sectioning along the axial direction 8-10.

Recently, quantitative phase imaging methods have drawn a broad attention due to their capability of imaging transparent objects with higher contrast in a quantitative manner 11. Especially, interferometry-based reflection microscopy (IRM) systems featured with widefield imaging capability for 2-D images, while keeping the lateral resolution comparable to that of the point-scanning microscopy, have been introduced, where the optical sectioning is produced by the coherence gating generated by the interferometric detection schemes 12-18. Due to the absence of a confocal pinhole or a nonlinear effect associated with highly condensed laser power, these methods can obtain the depth-selective 2-D image for an object with a single acquisition. The capability of complex field imaging, where the amplitude and phase information can be obtained simultaneously, enables IRM to measure physical properties of the target objects in a quantitative way. When combined with a transmission measurement, more information can be acquired on biological samples such 2 ACS Paragon Plus Environment

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as their physiological status

19.

Moreover, the label-free imaging capability, where no

exogenous labeling agent is required for imaging contrast, makes this technique more suitable for imaging of living samples.

Mainly, there have been two ways to achieve the optical sectioning in IRM. One is using time-gating produced by the short coherence length of a light source with broad spectrum 20-23. Only the light reflected from the region with a certain optical path length (OPL), which

is set by the position of a reference arm, can contribute to the signal, while light from other regions is rejected due to the short coherence length. Since the range of the contributing OPL is inversely proportional to the spectral bandwidth, employing a light source with broad frequency extent is a crucial requirement for achieving superior depth-selectivity in the time-gating approaches. For the time-gating, a reflection phase microscope employing a mode-locked Ti:sapphire laser was demonstrated in a wide-field 14 and also in a line-field mode 24. The former was featured with a fast acquisition rate up to 1 kHz based on a singleshot measurement using an off-axis detection configuration, and the latter the outstanding phase stability using the self phase-referencing method. However, due to the limited spectral bandwidth of the light source, the optical sectioning was not sufficient for distinguishing intercellular 3-D structures of living cells. A thermal light source was also utilized for reflection phase microscopy, where depth-selectivity of better than 1 μm was achieved by employing the wide spectral bandwidth from a halogen lamp 13, 15. This method successfully resolved the cellular structures in 3-D space and also the movement of filopodia of living cells 25, but was limited to slow acquisition rate involved with the phase-


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shifting detection scheme, as well as the insufficient light power associated with the nondirectionality of the light source.

Alternately, methods based on the spatial gating with a narrowband light source, where the optical sectioning is produced by the spatial coherence of light rather than the temporal coherence, were demonstrated. The Linnik-type interferometry was implemented with lasers with relatively long coherence lengths, of which spatial modes were scrambled randomly by passing the lasers through a rotating diffuser 12, or a multimode optical fiber 16.

Then, the time-varying speckle fields generated in 3-D space combined with the

interferometric detection achieved the optical sectioning originating from the decorrelation of the speckle fields along the axial direction. However, the off-axis detection configuration implemented by physically rotating the reference beam caused inevitable mismatch of the wavefront, which posed a limitation on the generation of the optical sectioning. Thus, the achievable depth-selectivity remained about 8 μm in both the cases.

We also have demonstrated reflection phase microscopy using dynamic speckle illumination (DSI) where the optical sectioning was either fully or partially determined by the spatial coherence of light

17, 26.

Unlike the previous implementations, the off-axis

detection port was constructed with no physical rotation of the reference beam by using a diffraction grating in the light path 17, 23. By accomplishing an almost perfect match of the wavefronts up to the maximum available numerical aperture (NA), a depth-selectivity of about 1 μm was achieved. However, the method based on the DSI was inefficient in utilizing the laser power, due to the limited use of the speckle field associated with its


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Gaussian-shaped angular distribution. Consequently, it was difficult to have enough signal from the samples with low reflectivity such as living eukaryotes in culture media even with a high-power light source.

In this article, we demonstrate reflection phase microscopy using a narrow-bandwidth plane wave, which is yet capable of depth-selective and wide-field imaging of living cells by a single-shot measurement. The depth selectivity is achieved by the successive accumulation of interferograms while synchronous scanning the angle of plane waves on a sample and a reference planes that are configured identically. The scanning covers the entire illumination NA of the system, resulting in narrow spatial gating along the axial direction. We theoretically addressed and numerically validated the origin of the depth selectivity by the SAI process. In addition, we experimentally implemented a reflection microscope based on this principle and achieved a depth-selectivity of about 0.76 μm. Since a plane wave is used for the illumination, our method can utilize all the laser power for the imaging purpose. With all these capability and performance of SAI, we demonstrated highresolution 3-D imaging of living cells in a culture medium and small worms under anesthesia, as well as high-precision measurements of fast dynamics occurring on cellular membranes.

Results Theory of SAI

The basic setup for implementing SAI is Linnik-type interferometry combined with an offaxis detection configuration by a grating (see Fig. 2a). Before discussing the experimental details, we first present in this section the theoretical framework for the working principle 5 ACS Paragon Plus Environment

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of SAI using the simplified system shown in Fig. 1a. Two plane waves from the same light source are introduced on the sample and the reference planes via two objective lenses, where the two arms are constructed identically in the Linnik-type configuration. To investigate the axial response of the system, a flat mirror is considered to be on the sample plane with an optical path difference (OPD) of z relative to the reference mirror. For taking the experimental situation, where a sample is immersed, into account, the space between the objective lens and the mirror in each arm is assumed to be filled with an immersion medium of refractive index 𝑛. Since the two arms are optically identical, the two beams have the same wavevector 𝑘 = (𝑘𝑥,𝑘𝑦,𝑘𝑧) = 𝑛𝑘0 associated with the illumination angle on both planes, where 𝑘0 = 2𝜋/𝜆0, where 𝜆0 is the center wavelength of the light source. The reflected light from the reference beam obtains an additional 𝑅

wavevector 𝑘 = (𝑘𝑅𝑥,0,0) along the x-axis from the grating, while the direction of the sample beam remains unaltered in the detection port. The separation of the two beams can be achieved by two linear polarizers located in the Fourier plane of which axes are perpendicular to each other. Thus, as shown in Fig. 1b, the two beams meet at a detector plane with the relative angle associated with 𝑘𝑅𝑥. According to this angle, a periodic fringe pattern is formed at the detector plane. For simplicity, we assume that the overall magnification from the sample plane to the image plane is unity. At the detector plane, the electric fields for the sample and the reference beams are given by𝑈𝑆(𝑥,𝑦,Δ𝑧;𝑘) = 𝐸𝑆(𝑥,𝑦)𝑒𝑖(𝑘𝑥𝑥 + 𝑘𝑦𝑦 + 2𝑘𝑧Δ𝑧)

and

𝑅

𝑈𝑅(𝑥,𝑦,0;𝑘) = 𝐸𝑅𝑒𝑖((𝑘𝑥 + 𝑘𝑥 )𝑥 + 𝑘𝑦𝑦),

respectively, aside from the common phase factor associated with the propagation through the detection port. 𝐸𝑆(𝑥,𝑦) = 𝐸𝑆 is the amplitude of the wave reflected from the sample


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mirror, and 𝐸𝑅 is the amplitude of the planar reference beam. The interferogram generated by the two beams with the specific wavevector can then be described by: 2

𝐼(𝑥,𝑦;𝑘) = |𝑈𝑆(𝑥,𝑦,Δ𝑧;𝑘) + 𝑈𝑅(𝑥,𝑦,0;𝑘)| .

(1)

While opening the detector within a single exposure, the wavevector for illumination, 𝑘, is scanned from 0 to 𝑘𝑚𝑎𝑥, corresponding to the maximum illumination angle supported by the objective lens. Figure 1b shows the situation occurring on the detector plane. During the scanning, the various interferograms are continuously accumulated on the detector. Thus, the cumulative interferogram can be obtained by integrating Eq. (1) over the scanning range as

𝑘

𝐼total(𝑥,𝑦) = ∫0𝑚𝑎𝑥𝐼(𝑥,𝑦;𝑘)𝑑𝑘𝑥𝑑𝑘𝑦. Interestingly, although the

integration is performed over the incoherent intensity signal, the result turns into a coherent accumulation of the electric field of the sample beam. This can be manifested by applying the Hilbert transformation to 𝐼𝑡𝑜𝑡𝑎𝑙(𝑥,𝑦). Thus the reflection intensity signal produced by the SAI can be calculated as:

S where,

1  2 2 12 2 1 2   cos 1 2 1  sin  1    , 4   

  2nzk0 ,

(2)

and   1  (NA n) 2 , with NA the numerical aperture of the

objective lens. See Supporting Information 1.1 for the details of Eq. (2).

Due to the dependence on ∆𝑧, S has a sharp peak around ∆𝑧 = 0, as shown in Fig. 1c. This can be understood with the accumulation process of the interferograms. When there is no path length difference (z = 0), the fringe patterns are stationary regardless of the scanning angle, thus the cumulative interferogram

I total ( x, y, z  0) maintains the


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interference contrast. However, when z is introduced, the fringe patterns have different lateral shifts depending on the illumination angle, resulting in loss of contrast while accumulating the interference

I total ( x, y, z  0) . With the specifications of the objective

lens (water immersion, NA=1) and the center wavelength of the laser, the full-width at half-maximum (FWHM) of S is expected to be about 660 nm. Since the coherence length of the light source is assumed to be sufficiently larger than the width of S, the depthselectivity is determined only by the SAI process. We also investigated the dependence of the central peak width on the NA of the objective lens. As expected, the reduction of NA causes the increase of the width of S as shown in Fig. 1c, and consequently worsens the depth-selectivity of SAI. It should be noted that we present the source coherence property as a line for zero NA since it corresponds to a usual interferometry with the stationary beam configuration.

Experimental setup

We experimentally implemented a reflection phase microscope based on the working principle described in the earlier section. Figure 2a shows a schematic of the experimental setup. A mode-locked Ti:sapphire laser (Coherent, Chameleon ultra II) with a center wavelength λ0 = 700 nm and spectral bandwidth ∆λ ≈ 7 nm was used as a light source. The collimated beam was split by a polarizing beam splitter (PBS) located between 4-f telescopes composed of a tube lens (TLi) and two objectives (OLS and OLR). A half-wave plate (HWP) was placed in front of the PBS to adjust the power balance between the sample and the reference arms by controlling the polarization axis of the input light. The redistribution of the light power was useful especially for measuring reflection signals from 8 ACS Paragon Plus Environment

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biological samples with very low reflectivity. For the identical configuration of the interferometry, two objective lenses with the same specification were used (water immersion, 1.0 NA, 60×, Olympus). A blank mirror was placed at the focal planes of the reference arm. For the sample arm, either a blank mirror for system alignment, or a sample for measurement, was used. A quarter-wave plate (QWP) was placed in each arm of which the retardation axis was set to make the polarization of the light way out perpendicular to that of the light way in. Thus, the direction of the returning light was tuned to the output port of the PBS. The illumination angle of the laser beam was controlled by a dual-axis galvanometer scanning mirror (GM, Saturn 1B, Pangolin Laser system, Inc.) positioned in one image plane conjugate to the sample plane. The identical optical configuration in the two interferometric arms allows the illumination beams to scan an object and the reference mirror with the same incident angles synchronously.

A diffraction grating (Ronchi Rulings, Edmund Optics) was placed in the conjugate image plane to implement the off-axis configuration without rotation of the wavefront 17, 23.

The diffracted light was delivered to a camera (Flea3, CMOS, PointGrey) through a 4-f

telescope (L1-L2), during which multiple spots associated with the diffraction orders were spatially separated in the Fourier plane of the 4-f telescope. For imaging purpose, only the 0-th and 1-st diffraction orders were selected, and all other orders were blocked by a spatial filter (not shown in Fig. 2). To separate the sample and the reference beams one from the other, two linear polarizers, of which the polarization axes were set to be perpendicular to each other, were placed right behind the spatial filter at the Fourier plane. Subsequently, two QWPs were used, one with +45 degrees, and the other with -45 degrees. Consequently,


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the two beams have circular polarizations with the same handedness, resulting in the interference fringes with the maximum contrast.

The generated interferograms were captured by the camera. While opening a single exposure, the GM scanned the illumination beam, such that the focused spots in the back apertures of the objectives uniformly addressed the entire apertures in a spiral shape. Thus, the scanning covered the whole NA of the objectives with uniform density. From the system characterization, we determined the number of scanning angles N larger than that predicted by the simulation (see Supporting Information 1.2). Typically, N = 200 was used for acquiring a single object image. The camera was triggered with the GM scanning in such a way that multiple angle-dependent interferograms were accumulated at the camera sensor during one cycle of the whole scanning process.

Investigation of depth-selectivity

First, we investigated the depth-selectivity of our system by measuring its axial response. For this experiment, a flat mirror was used as a sample, and its image was measured by SAI. We repeated the measurements at different positions of the sample mirror, while holding the reference mirror position. The sample mirror was moved 20 μm in total with an interval of 100 nm by a motorized stage. The upper inset in Fig. 2b shows the accumulated interferogram when there is no OPD between the two arms. Since all the interferograms were accumulated with no phase shift, the fringe pattern shows the maximum contrast. In contrast, when the sample mirror was moved by 1 m, a weak fringe pattern was observed, as shown in the lower inset in Fig. 2b. Unlike the case of no OPD,


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the fringe contrast significantly reduced because all the interferograms were added with different angle-dependent phase shifts. All the interference raw images of the sample mirror were processed into complex field images using the standard Hilbert transform method

27,

and the intensity maps were obtained by taking the absolute square of their

complex maps. Then, the averaged value, which was extracted from a 68.5 μm2 area in each intensity map, was measured as a function of the OPD of the sample mirror.

Figure 2b shows the SAI’s characteristic signal response S along the axial direction when using the full NA of the objective lens for illumination. As expected theoretically and numerically, a sharp peak at the zero OPD was observed. In order to investigate the effect of sampling number N for the angular scanning in the back aperture, we varied N as 10, 50, 100, and 200. For a low sampling number with N = 10, a complicated response with many irregular side lobes was observed. When using a higher sampling number such as N = 100 and 200, the side lobes were effectively suppressed, and the response curve became smooth with a well-defined peak at the center. For various N, the widths of the central peaks remained almost unchanged. Regardless of N, the FWHMs of the axial responses were measured to be about 0.76 ± 0.02 m. After combining the SAI configuration, the gating capability is enhanced by a factor of 55 times compared to the inherent coherence response of the laser itself (see Supporting Information 2.1). But, this value is slightly larger than that expected in theoretical and numerical investigation. This deviation was attributed to the slight mismatch of the two wavefronts at the detector plane, and uncontrollable system vibration during the accumulation process of interferograms.


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Although the numerical simulation verified that about N = 50 was sufficient for suppressing the unwanted side lobes in the axial response, the results in Fig. 2b show that more than N = 100 is required for SAI to obtain well-defined depth-selectivity with reasonable quality. The reasons for this discrepancy are thought to be the digitized nature of the camera acquisition and the angle-dependent signal variation, which were not accounted for in the simulation. Although numerous angles of illumination are essential for well-defined depth selectivity, choosing too large an N sacrifices the acquisition speed. With this trade-off, we fixed N = 200 for the well-shaped depth-selectivity minimizing the compromise of the acquisition speed, and this was used for all the remaining experiments hereafter.

We also varied the coverage of the angular scanning in the back focal plane of the objective lenses. This effectively adjusted the illumination NA of the system. Figure 2c shows the measured axial responses for the several illumination NAs. Similar to the numerical results presented in Fig. S1d in Supporting Information 1.2, the axial response broadened and thus the depth-selectivity became worse for smaller angular coverage. This conforms to the theoretical and numerical predictions that covering the whole back aperture of the objective is essential to achieve the highest depth-selectivity. Note that, in Fig. 2c, the case of zero NA shows only the effect of time-gating measured with the Ti:sapphire laser (see Supporting Information 2.1).

With the fixed N = 200, the acquisition speed was determined by the scanning rate of the GM. We investigated the shape of the axial response depending on the imaging speed.


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The camera exposure was varied, such that the acquisition rate became 10, 50, and 100 fps, and the scanning speed of the GM was adjusted as 0.2, 1, and 2 kHz accordingly. The shape of the axial response was slightly corrupted at a higher frame rate while keeping the central width unchanged, as shown in Fig. 2d. Since this degradation in the imaging quality is negligible, we used 100 fps for the remaining experiments. This limitation is set by both the camera acquisition rate and the GM scanning rate (see Online Methods). If a higher speed camera and faster scanning equipment is used, the acquisition speed can be improved further.

3-D imaging of living cells

We applied the SAI method to the imaging of living cells. Breast cancer cells (MDA-MB231) were prepared on a collagen layer on a glass slide, to minimize the specular reflection from the cell-glass interface. A cluster of multiple cells with different heights was chosen for imaging, as schematically shown in Fig. 3a. Multiple images were taken while moving the sample in the downward direction so that the focal plane where the optical sectioning occurred was moved in the upward direction. The overall scanning range was 18 μm, with a step size of 200 nm. Thus, 90 images for the cells were acquired in total at different depths.

Figures 3b and 3c present representative images taken at two different depths. Figure 3b shows the intensity image for the cell cluster when the focus lay on the plane indicated by the red dashed line in Fig. 3a. In that figure, the top surface of the smallest cell was in focus. Since the cell membrane was located right in the imaging range, the top portion of the smallest cell was seen as denoted by the yellow circle in Fig. 3b. In contrast, when moving the sample by 1 μm in the downward direction, the focal plane was raised above 13 ACS Paragon Plus Environment

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the cell membrane by 1 μm and the cell was out of the imaging range. Consequently, the signal from the surface was rejected by the SAI process and thus it was almost invisible in Fig. 3b.

We constructed a 3-D map for the cell cluster using the whole set of depth-selective images. Figure 3d shows the 3-D map generated over the volumetric ranges of 85.3 µm × 106.6 µm × 18 µm. The structures and physical quantities, such as locations, sizes, heights, and intercellular distributions of the individual cells, are clearly resolved in the 3-D space. Supplementary Movie 1 shows the 3-D rendered image for those cells.

We compared the laser power required for SAI to obtain the reflection images similar to those acquired by the DSI configuration. For the images shown in Fig. 3, the input powers for SAI and DSI were 1.19 mW and 51.1 mW, respectively, meaning that SAI used only 2.3 % of the input power that the DSI configuration needed for acquiring the reflection images with the similar level of signal strength. This difference in power efficiency of more than 43 times is due to the lossless nature of SAI in light utilization, which is critical especially for biological samples with substantially low reflectivity such as living eukaryotes.

Extended imaging for 3-D visualization of a living worm

We also applied SAI for the imaging of a sample of which the dimension is beyond the cellular level. A living Caenorhabditis elegans (C. elegans) was prepared in a state of immobilization (see Online Methods). As done with the living cells in Fig. 3d, whole-depth images of the C. elegans were taken, while varying the position of the sample plane. In 14 ACS Paragon Plus Environment

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order to enhance the image quality, 10 images taken at the same depth were averaged to produce a single depth-selective image. Since the C. elegans was larger than the lateral field of view of the system, we repeated the volumetric imaging at several positions to cover the entire dimension of the C. elegans.

Figures 4a–4c show the representative depth-selective images taken for the head part, denoted by the yellow box in the inset in Fig. 4d. On the outmost surface of the worm, at a depth of 36.4 μm, a striped-pattern, according to the cuticle structure, can be seen clearly in Fig. 4a. When going deeper into the worm, some characteristic structures associated with internal organs could be observed. At a depth of 27.2 μm in Fig. 4b, a structure with relatively strong reflection compared to the neighboring region was observed. The round shape of the structure shows the typical feature of a pharynx of the worm. At 14 μm in Fig. 4c, a long streak-like structure through the body was observed. This seems to be a part of the esophagus of the worm 28.

We produced a z-stack image of the entire body of the worm by stitching the z-stack images taken at five different imaging sites, as shown in Fig. 4d. The overall posture and structure of the worm can be observed. We also mosaicked the five volumetric images to form an extended 3-D image for the whole C. elegans, as shown in the inset in Fig. 4d. The overall dimension for this extended 3-D image is about 281 µm × 191 µm × 40 μm. The curved formation of the worm and the fine details for its internal structures are clearly visualized in 3-D space. (see Supplementary Movie 2 showing the 3-D volume data rendered by AMIRATM.)


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Measurement of the membrane fluctuation of a living cell

We used SAI to measure fast dynamic motions in the plasma membranes of living cells. Fine membrane fluctuation was measured in phase variations over time and then converted into height changes. As a target sample, lung cancer cells (H460, human lung-cancer epithelial cells) in a culture medium were used. The cells were prepared on a polydimethylsiloxane (PDMS) layer about 3 mm thick, placed on a petri dish to minimize the specular reflection from the dish surface. To measure the reflection signal from the plasma membrane of the cell, the focus was placed on the outmost top surface of the cell. At that position, we acquired 1,000 images at the speed of 100 fps.

Figure 5a shows the intensity distribution reflected from the top plasma membrane of the cell. Due to the optical sectioning, only a small portion of the top surface is visible. Figures 5b–5d show representative reflection phase distributions obtained at different times (t = 4, 7, and 10 s, in order). Over the measurable area of the plasma membrane, we chose only the region where the temporal evolution of the reflection phase remained continuous during the entire measurement time. Other regions that showed broken phase variations associated with sudden jumps or unreasonably fast oscillations, as well as empty background areas revealing random scattered phases, were removed and are represented in black in Fig. 5. Around the center region as indicated with the red arrow in Fig. 5c, a 5 pixel × 5 pixel area corresponding to the diffraction-limited spot in our system was chosen, and the average phase values over the pixels were traced over time. The result is presented in Fig. 5i with the red line. The root mean square (RMS) magnitude of the membrane fluctuation was 385.4 mrad, which corresponds to 16.14 nm in length. The black line in Fig. 5i shows phase fluctuation measured at blank area where there was no sample. The 16 ACS Paragon Plus Environment

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RMS magnitude was measured as 24.6 mrad, representing our characteristic measurement stability of 1.03 nm.

Next, we treated the same cell with a fixation medium of 4% paraformaldehyde (PFA). After the PFA treatment, the intensity distribution from the top membrane of the same cell is shown in Fig. 5e. The difference in the shape was attributed to the slight morphological change during the PFA treatment. After the treatment, the membrane fluctuation was observed in the same position, as indicated by the blue arrow in Fig. 5g. The result is presented in Fig. 5i by the blue line. The RMS fluctuation in phase in turn decreased to 100.8 mrad (4.22 nm), which is about a fourth of that before the PFA treatment. Figures 5j and 5k show 2-D maps for RMS magnitudes of the fluctuations of the living cell and the treated cell, respectively. It can be seen that the membrane motion after the treatment was reduced noticeably.

Discussion and Conclusion The optical sectioning of SAI is generated by the synchronous scanning of the illumination angles of the two plane waves on the sample and the reference planes. The angle-dependent phase shifts on the returning wave from the sample, which are caused by the OPD relative to the reference plane, reduce the cumulative fringe contrast and thus rapidly wash out the out-of-focus contribution. The degree of this suppression, which determines the depth-selectivity of the system, strongly depends on the uniformity of the illumination coverage in the back focal (BF) plane of the objective lens but is independent of the spectral bandwidth of the light source. When comparing with the former


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implementation using the decorrelation of the time-varying speckle fields, SAI has certain advantages in utilizing the illumination light because it employs a scanning plane wave. With the complete control of the plane wave, SAI can perform the uniform scanning of the light over the entire BF plane of the objective lens with no loss of its power, which guarantees the maximal depth-selectivity with the minimal light energy.

The main purpose of our method is the high-speed and high-precision acquisition of the depth-selective images at a specific target depth of a sample. The SAI process provides the optical sectioning sufficient for resolving the 3-D details of cellular sturcutres. Thus the 3D reflectance image similar to that obtained by the confocal reflectance microscopy can be constructed by collecting multiple depth-selective images taken while moving the axial position of the sample. Since only the reflected signal is mapped in the space, this method cannot visualize the physical property other than the reflectance distribution. In this sense, our method has a certain limitation in 3-D imaging when compared with the tomographic techniques in a transmission geometry

29-33,

which can reveal the 3-D refractive index

distribution of a sample with the use of the cumulative phases experienced by light passing through the sample under various illumination states. However, since our method acquires a single-depth image with a single-shot measurement and also with the axial motion sensitivity of about 1 nm, it has a strong advantage for the observation of fast and fine dynamics of the layered structures such as cellular membranes. In addition, its reflection configuration allows to measure the sample from the outmost surface up to the penetration depth regardless of the sample’s thickness. This makes our method have more potentials for actual biomedical applications compared to the techniques in the transmission geometry.


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In conclusion, we have developed a method for wide-field reflection phase microscopy using a narrow-band plane wave, which is capable of depth-selective imaging of an object with a single-shot measurement. We theoretically showed that the SAI process forms the depth-selectivity of about 0.66 μm under our conditions, and numerically investigated the optimal parameters for achieving this optical sectioning with the consideration of the discrete sampling situation. We experimentally implemented a reflection phase microscope working with the SAI and then demonstrated the depth-selectivity of 0.76 μm, which is fairly close to the theoretical and numerical expectations. We successfully acquired the depth-selective images of living cells in a culture medium and constructed the 3-D map of the cells from the multiple acquired images. In addition, by stitching several 3-D volumetric images taken at multiple sites, we could also generate the extended 3-D image of a C. elegans with dimensions far beyond the actual field of view of the microscope. Since our scheme employs a plane wave, it is free from the significant waste of light, which is inevitable in other imaging methods based on the varying speckle fields. With this advantage, SAI could obtain the reflection images of biological samples with the same level of signal strength using the minimal light power.

With all these performances and advantages, our method will facilitate nanometer-scale surface profilometry

34-37,

precision measurements for the fast dynamics associated with

cellular membrane motions

38-40,

and investigation of nanomechanical properties

41-44

of

individual single cells in a non-contact manner. This will also be useful for nondestructive


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3-D observation and topological measurements of biological samples, such as cells and small animals, as well as non-biological specimens.

Methods Determination of the acquisition speed by the GM scanning rate

Since the entire back apertures of the objective lenses should be covered by the scanning plane wave within a single exposure, the acquisition speed is limited by not only the camera speed, but also the GM scanning rate. The maximum sinusoidal scanning frequency of the galvanometer mirror used in the setup is 2 kHz at about ± 5° scanning angle. When reaching the edge of the back aperture of the objective lens, the illumination angle formed at the sample plane is about ± 48.7°. Considering the overall magnification from the front conjugate plane where the scanning mirror is placed and the sample plane, the maximum scanning angle is ± 4.32°. In addition, to achieve N = 200 in a spiral shape fitted in the back aperture of the objective, as was verified by the numerical simulation, about 10 sinusoids are required. Under these operating conditions, we can achieve the acquisition speed of 100 fps, including a 50% duty cycle, between the forward and the backward scans.

Preparation of living cells on a collagen layer

The collagen gel layer was prepared by mixing collagen type I (rat tail, 3.5 mg/mL; Corning) with 10× phosphate buffer saline (Gibco), deionized water, and 0.5 N NaOH. The collagen solution was then plated on the 35  petri dish (DB Falcon) with a thickness of about 3 mm at a concentration of 2.5 mg/mL. For gelation, the prepared petri dish was stored in an incubator at 37°C for 40 minutes.


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Human Breast cancer cells (MDA-MB-231) were cultured and maintained on the prepared collagen gel layer at a concentration of about 100 cells/mm2 in Roswell Park Memorial Institute (RPMI) 1640 medium (Gibco), supplemented with 10% fetal bovine serum (Gibco), and 1% (v/v) penicillin and streptomycin (Gibco). After the seeding, the whole petri dish was stored in an incubator at 37°C with 5% CO2 for 4 hours for cell stabilization.

Preparation of a living C. elegans

The C. elegans used in our experiments was an N2 strain, purchased from Caenorhabditis genetic center (CGC, Univ. of Minnesota, USA). Before the experiments, worms were cultivated on the nematode growth media (NGM) plate (17 g Bacto agar, 2.5 g Bacto peptone, 3 g NaCl, 1 ml 1 M CaCl2, 1 ml 1 M MgSO4, 25 ml 1 M KPO4 buffer, and 1 ml 5 mg/ml cholesterol in ethanol to 975 ml H2O) with a lawn of Escherichia coli OP50 (CGC, Univ. of Minnesota, USA) as a food source. For experimental observation, we took a small block of an NGM plate 10 mm × 5 mm × 5 mm, where several worms (L2 stage, body length ~400 μm) were freely crawling on its surface. Then these worms were immobilized by treatment with a 100-mM sodium azide (NaN3, Sigma, USA) solution for 20 minutes 45

and prepared on the imaging plate for SAI measurement.

Associated Content Supporting Information Supporting Information 1: Theoretical frame work and numerical simulation for SAI Supporting Information 2: Generation of optical sectioning by SAI


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Supporting Information 3: The detailed scheme for experiment setup and comparison of spatial resolution and power consumption between SAI and DSI Supporting Information 4: conversion accuracy of phase to height Supplementary Movie 1: 3-D rendered image of the cells Supplementary Movie 2: Extended 3-D image obtained by making a mosaic of five volumetric data.

Author Information Corresponding Authors

*E-mail: [email protected] *E-mail: [email protected] ORCID

Min Gyu Hyeon: 0000-0002-2907-3444 Taeseok Daniel Yang: 0000-0001-6452-7504 Young Guk Kang: 0000-0002-5572-7544 Beop-Min Kim: 0000-0002-1056-5078 Youngwoon Choi: 0000-0001-8492-8225 Author contributions

M.G.H., B.-M.K., and Y.C. conceived the idea and designed the experiment. M.G.H. and Y.C. developed the theory and performed the numerical simulation. M.G.H. and T.D.Y. carried out the measurements and analyzed the data. T.D.Y., K.P., and Y.G.K. prepared the collagen gel layer and the cells, and J.-S.P. prepared the C. elegans. M.G.H., T.D.Y., J.-s.P., and Y.C. prepared the manuscript, and all authors contributed to finalizing the manuscript. Notes

The authors declare no competing financial interest. 22 ACS Paragon Plus Environment

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Acknowledgements This work was supported by the National Research Foundation of Korea (NRF) funded by the

Ministry

of

Science,

ICT

&

Future

Planning

(2017R1C1B2010262,

2017R1A6A3A11031083). It was also supported by IBS-R023-D1 for J.-S.P, and the Korea Health Technology R&D Project through the Korea Health Industry Development Institute (KHIDI), funded by the Ministry of Health & Welfare, Republic of Korea (HI14C3477), and a grant from Korea University.

References

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(28) Zhang, D.; Li, C.; Zhang, C.; Slipchenko, M. N.; Eakins, G.; Cheng, J.-X., Depthresolved mid-infrared photothermal imaging of living cells and organisms with submicrometer spatial resolution. Sci. Adv. 2016, 2 (9), e1600521. (29) Choi, W.; Fang-Yen, C.; Badizadegan, K.; Oh, S.; Lue, N.; Dasari, R. R.; Feld, M. S., Tomographic phase microscopy. Nat. Methods 2007, 4, 717. (30) Hosseini, P.; Sung, Y.; Choi, Y.; Lue, N.; Yaqoob, Z.; So, P., Scanning color optical tomography (SCOT). Opt. Express 2015, 23 (15), 19752-19762. (31) Park, Y.; Diez-Silva, M.; Popescu, G.; Lykotrafitis, G.; Choi, W.; Feld, M. S.; Suresh, S., Refractive index maps and membrane dynamics of human red blood cells parasitized by Plasmodium falciparum. Proc. Natl. Acad. Sci. U.S.A. 2008, 105, 1373013735. (32) Kim, K.; Yaqoob, Z.; Lee, K.; Kang, J. W.; Choi, Y.; Hosseini, P.; So, P. T. C.; Park, Y., Diffraction optical tomography using a quantitative phase imaging unit. Opt. Lett. 2014, 39 (24), 6935-6938. (33) Kim, T.; Zhou, R.; Mir, M.; Babacan, S. D.; Carney, P. S.; Goddard, L. L.; Popescu, G., White-light diffraction tomography of unlabelled live cells. Nat. Photon. 2014, 8, 256. (34) Radmacher, M.; Tillmann, R. W.; Fritz, M.; Gaub, H. E., From Molecules to Cells - Imaging Soft Samples with the Atomic Force Microscope. Science 1992, 257 (5078), 1900-1905. (35) Parpura, V.; Haydon, P. G.; Henderson, E., 3-Dimensional Imaging of Living Neurons and Glia with the Atomic Force Microscope. J Cell Sci 1993, 104, 427-432. (36) Kasas, S.; Ikai, A., A Method for Anchoring Round Shaped Cells for Atomic-Force Microscope Imaging. Biophys J 1995, 68 (5), 1678-1680. (37) Rotsch, C.; Braet, F.; Wisse, E.; Radmacher, M., AFM imaging and elasticity measurements on living rat liver macrophages. Cell Biol Int 1997, 21 (11), 685-696. (38) Schneider, S. W.; Sritharan, K. C.; Geibel, J. P.; Oberleithner, H.; Jena, B. P., Surface dynamics in living acinar cells imaged by atomic force microscopy: Identification of plasma membrane structures involved in exocytosis. Proc. Natl. Acad. Sci. U.S.A. 1997, 94, 316-321. (39) Dulebo, A.; Preiner, J.; Kienberger, F.; Kada, G.; Rankl, C.; Chtcheglova, L.; Lamprecht, C.; Kaftan, D.; Hinterdorfer, P., Second harmonic atomic force microscopy imaging of live and fixed mammalian cells. Ultramicroscopy 2009, 109 (8), 1056-1060. (40) Preiner, J.; Tang, J. L.; Pastushenko, V.; Hinterdorfer, P., Higher harmonic atomic force microscopy: Imaging of biological membranes in liquid. Phys Rev Lett 2007, 99 (4). (41) Cross, S. E.; Jin, Y. S.; Rao, J.; Gimzewski, J. K., Nanomechanical analysis of cells from cancer patients. Nat Nanotechnol 2007, 2 (12), 780-783. (42) Iyer, S.; Gaikwad, R. M.; Subba-Rao, V.; Woodworth, C. D.; Sokolov, I., Atomic force microscopy detects differences in the surface brush of normal and cancerous cells. Nat Nanotechnol 2009, 4 (6), 389-393. (43) Raman, A.; Trigueros, S.; Cartagena, A.; Stevenson, A. P. Z.; Susilo, M.; Nauman, E.; Contera, S. A., Mapping nanomechanical properties of live cells using multi-harmonic atomic force microscopy. Nat Nanotechnol 2011, 6 (12), 809-814. (44) Gavara, N.; Chadwick, R. S., Determination of the elastic moduli of thin samples and adherent cells using conical atomic force microscope tips. Nat Nanotechnol 2012, 7 (11), 733-736.


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(45) Franks, C. J.; Holden-Dye, L.; Bull, K.; Luedtke, S.; Walker, R. J., Anatomy, physiology and pharmacology of Caenorhabditis elegans pharynx: a model to define gene function in a simple neural system. Invertebr. Neurosci. 2006, 6 (3), 105-122.


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Figure 1

Fig. 1. Conceptual Schematic of SAI. a, Synchronous angular scanning of a plane wave on both the sample and the reference planes. The two beams are steered such that they have the same angles on their focal planes. Immersion media with refractive index n are introduced on both the planes. The sample plane is shifted by ∆z from that of the reference plane. The illumination angle θ is varied from 0 to max covering all the conical volume subtended by max. OLS and OLR: objective lens for the sample and the reference arms; max: maximum illumination angle supported by the objective lens; BF: back focal plane of the objective lens; M: mirror. b, Successive accumulation of interferogram. When z = 0, all interference patterns are stationary resulting in the maximum contrast in the cumulative interferogram. If z ≠ 0, the interference patterns are shifted depending on the illumination angle, thus the cumulative interferogram loses the contrast. c, Axial response S of SAI as a function of ∆z calculated with various 𝜃0 corresponding to 0.0, 0.4, 0.6, 0.8, and 1.0 NA, of which filling rations of the back focal plane of the objectives are 0%, 16%, 36%, 64%, and 100% in order. FWHM of S is 0.66 μm when covering the whole NA of the objective lens.


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Figure 2

Fig. 2. Experimental setup and the depth-selectivity. a, Setup schematics. Laser: mode-locked Ti:sapphire;

GM: dual-axis galvanometer mirror; M: mirror; L1 and L2: lenses; TLi and TLo: tube

lenses for input and output ports; PBS: polarizing beam splitter; HWP: half-wave plate; QWP: quarter-wave plate; OLR and OLS, objective lenses for reference and sample arms; G: grating; LP: linear polarizer. b, Axial response S as a function ∆z with various numbers of scanning angles N. Insets: cumulative interferograms in single camera exposures with ∆z = 0 (top) and ∆z = 1 μm (bottom). c, Axial responses S with various illumination NAs which were adjusted by controlling the angular coverage of the GM scanning. d, Axial responses S depending on the acquisition speed which was determined by the GM scanning rate. For SAI, N = 200 was used for measurements in c and d.


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Figure 3

Fig. 3. 3-D structural imaging of living cells. a, Schematic diagram for the formation of the cells (side view). b, Depth-selective intensity image of the cells with the focal plane indicated by the red dashed line in a. The top surface of the smallest cell can be seen, as denoted by a yellow circle. Scale bar: 20 m. c, Same as b, but with the focal plane shifted by 1 μm, as denoted by the black dashed line in a. The smallest cell disappeared due to the optical sectioning. d, 3-D volume rendered image for the cells.


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Figure 4

Fig. 4. Extended image of a C. elegans taken by SAI. Depth-selective images obtained at different depths of a 36.4 μm, b 27.2 μm, and c 14 μm from the bottom plane of volume 1 in d. Scale bar: 20 μm. Insets: magnified images for the regions in the small white boxes. Scale bars: 5 μm. d, zstack image produced by stitching image taken at five different imaging sites. Inset: extended 3-D image obtained by making a mosaic of five volumetric data. Scale bar: 50 μm. The whole dimension of the extended volume is 191 µm × 281 µm × 40 μm.


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Figure 5

Fig. 5. Dynamics of a cell membrane. a, Distribution of reflection intensity from the top plasma membrane of the living cell. Scale bar: 5 m. b–d Reflection phase distributions on the top membrane of the living cell measured at time t = 4, 7, and 10 s, in order. e, Same as a, but after the treatment of 4 % PFA. Scale bar: 5 m. f–h Same as b–d after the PFA treatment. i, Time traces of the phase fluctuations measured at point A indicated by the red arrow in c (red), at point B indicated by the blue arrow in g (blue), and a blank area (black). Distributions of RMS magnitudes of the fluctuations j before and k after the treatment. Color bar in a: normalized intensity (A.U.), and color bars in b–d, f–h, j, and k: phase in radian.


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For Table of Contents Use Only

Reflection phase microscopy by successive accumulation of interferograms Min Gyu Hyeon, Taeseok Daniel Yang, Jin-Sung Park, Kwanjun Park, Yong Guk Kang, Beop-Min Kim, and Youngwoon Choi

The TOC graphic schematically represents the process of the successive accumulation of interferograms (SAI) with the synchronous scanning of an illumination beam in the reflection phase microscope. The interference contrast at the detector strongly depends on the path length difference between the two arms, thus the SAI process provides an optical sectioning sufficient for resolving multiple layered structures of biological samples, such as living eukaryotes and small animals. Since the development is capable of wide-field acquisition, it is suitable for measuring fast and fine dynamics occurring in the layered structures of biological specimens.


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Reflection phase microscopy by successive accumulation of

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