2.1.

Experimental Setup of the Microscope and Device

Observations were performed with an upright microscope (Eclipse FN1, Nikon Co.) and a scanner (A1R MP+, Nikon Co.) for excitation light with wavelengths 488 and 950 nm. In this study, the confocal aperture was completely open and the 13-mm-thick device was inserted between a water-immersion objective lens (CFI Apo LWD $25\times \mathrm{W}$, NA1.1, Nikon Co.) and the microscope revolver. A pulse amplitude modulation controller was used to control the performance of the device [Fig. 1(a)]. Figures 1(b)–1(d) illustrate the principle of the spherical aberration correction by the LCC. Figure 1(b) shows an example of a spherical aberration distribution and Fig. 1(c) shows a schematic of the six-level quantized phase distribution, which corrects the spherical aberration. Figure 1(d) shows the segmented annular electrodes and lead-out electrodes for applying voltage to the liquid-crystal layer to generate the discretized phase distribution in the LCC. The diameter of the outermost peripheral annular electrode was 17.6 mm, which was almost equal to the calculated pupil diameter of the objective lens, according to the manufacturer’s information. In this study, we first applied a voltage determined from the numerical calculation. Next, by repeatedly changing the applied voltage slightly, we determined the appropriate voltage required to achieve the maximized fluorescence signal of the object.²⁷

Fig. 1

Experimental setup: (a) configuration of the laser scanning microscope with an adaptive optics device containing liquid crystal cells (LCCs); schematic of (b) spherical aberration and (c) phase distribution of the LCC; (d) schematic of the transparent electrodes in the LCC. (e) Calculated Strehl ratio after correcting spherical aberrations by a quantized phase distribution. (f) The actual phase distribution of the device and the phase profile of the center of the phase distribution. (g) Transmittance of the LCC.

After the correction, the discretized phase distribution in the LCC caused a residual aberration, which can be reduced by increasing the number of rings in the phase distribution in the LCC. However, the number of annular electrodes was limited by the production process and could not be further increased. In addition, even if the number of annular electrodes was increased, the corresponding increase in the number of lead-out electrodes would cause a deterioration of the spatial homogeneity and, thus, the optical performance. Therefore, we used numerical calculations to evaluate and optimize the ring number to obtain a maximum effect of correcting the spherical aberration. The numerical calculations were performed based on the Fourier transforming property of lenses. A fast Fourier transform was applied to the phase distribution of the residual aberration on the pupil plane of the objective lens, which resulted in an amplitude distribution on the focus plane. However, the device was not placed in the pupil position of the objective lens during actual implementation. The influence of the displacement of the device from the pupil position along the $z$-axis is considered in the next section. For an NA1.1 water-immersion objective lens and a sample with a refractive index of 1.38, the amplitude of the induced spherical aberration at a depth of 0.7 mm was calculated as 3600 nm. This value was equal to that of the phase modulation, which can be generated by using three LCCs. The phase distribution quantized in 48 or more levels resulted in a Strehl ratio greater than 0.9 [Fig. 1(e)]. Hence, we used three LCCs to generate a 48-level quantized phase distribution. The LCCs had 31 segmented annular electrodes with linewidths 1.3–0.043 mm. The device achieved a phase modulation of 3600 nm. The phase distribution generated in the device was measured by laser interferometers (Zygo PTI 250, Zygo Corporation) and we confirmed that its form agreed very well with that of the calculated distribution [Fig. 1(f)].

Figure 1(g) shows the transmittance of the LCC. For the wavelength range 800–1100 nm, which was used for excitation in TPLSM, the transmittance was 95% per cell. For the wavelength range 460–600 nm, which was used for fluorescence measurements, the transmittance was 92%–95% per cell. Since we used three LCCs in this study, the transmittance of the device was approximately 88% and 78%–85% in the wavelength ranges 800–1100 and 460–600 nm, respectively.

2.2.

Numerical Calculations Using the Objective Lens System

The point spread function (PSF) of the 488-nm collimated light beam was numerically calculated (CODEV, Optical Research Associates) assuming that the light was focused through a continuous phase distribution and a $25\times /\mathrm{NA}1.1$ water-dipping objective lens (JP2011-75982A). The numerical calculation was performed based on the beam propagation method. In this calculation, the continuous phase distribution was placed in a plane perpendicular to the optical axis, 10 mm away from the front of the objective lens along the $z$-axis (where the device was actually placed in the experiment). To evaluate the effect of the spatially discretized phase distribution of the device with a 48-level quantized phase distribution, the reduction rate of the Strehl ratio was also obtained in the numerical calculation, as mentioned in Sec. 2.1. Therefore, we were able to estimate the influences of both the device’s position along the $z$-axis and the discretization of the phase distribution.

2.3.

Sample Preparation and Analysis

Fluorescent yellow-green beads with a diameter of $0.2\mu \mathrm{m}$ were mixed into agarose gel with a refractive index of 1.38. The refractive index of agarose gel was adjusted using sucrose solutions. The mixture was poured into a glass base dish (cover glass thickness of 0.16–0.19 mm). Under different observation conditions, we observed several beads within the volume of a $20\text{-}\mu \mathrm{m}$ cube located at the center of the field of view. The fluorescence intensity profiles across the intensity center along the $x$-axis were fitted with a Gaussian function, and the peak intensities and full width at half maximum values along the $x$-axis ($\mathrm{FWHM}x$ values) were obtained for each image. Thus, the average peak intensity and the average $\mathrm{FWHM}x$ value of the images were obtained, along with the average intensity profile of those images. Similarly, fluorescence intensity profiles across the intensity center along the $z$-axis were fitted with a Breit-Wigner-Fano line shape for each image, and the average FWHM value along the $z$-axis ($\mathrm{FWHM}z$ value) and the average intensity profile of those images were obtained.

Previously, 2,2’-thiodiethanol (TDE) solutions were shown to quickly render fixed mouse brain tissues optically transparent.^7,8 TDE-treated fixed mouse brain slices exhibited a higher penetration depth in the visible and the near-infrared regions, and a stronger fluorescence signal was emitted from the deep layers. In this study, we used 51% TDE solutions to clear the 0.95-mm-thick fixed brain slices of YFP expressing mouse (Thy1-YFP-H mouse²⁸) and to adjust the refractive index to 1.42. The TDE-treated slices were immersed in a TDE solution of the same concentration in the mount, where the distance between the cover slip and the slide glass was 1 mm. To estimate the spatial resolution in a fixed brain slice, we utilized dendritic spines about $1\mu \mathrm{m}$ in size as a ruler in the cortical layer V neurons. The intensity profiles of a single dendritic spine along the $x$- and $z$-axes were fitted with a Gaussian function and a Breit-Wigner-Fano line shape, respectively, to obtain the peak intensities and the FWHM values along the $x$- and $z$-axes.

All animal experiments were performed in accordance with the National University Corporation Hokkaido University Regulations on Animal Experimentation and the Guidelines for Proper Conduct of Animal Experiments (Science Council of Japan). The protocol was approved by the Institutional Animal Care and Use Committee of National University Corporation Hokkaido University (Permit No. 10-0119).

3.1.

Numerically Calculated Point Spread Function at 488-nm Wavelength

First, in order to estimate the correction effect of the device, we numerically calculated the PSFs at the surface of the cover glass ($z=0\mathrm{mm}$) and at depths of 0.3 and 0.6 mm in a sample with a refractive index of 1.38 (Fig. 2). The spatial resolution and the signal of the PSFs degraded in proportion to the depth [Figs. 2(a)–2(d), 2(g), 2(h)]. At a depth of 0.3 mm from the surface of the sample ($z=0.3\mathrm{mm}$), the peak intensity of the PSF dropped to 13% of the peak intensity at $z=0\mathrm{mm}$ ($I_0$). Additionally, the $\mathrm{FWHM}x$ value degraded to $0.32\mu \mathrm{m}$ from $0.22\mu \mathrm{m}$ and the $\mathrm{FWHM}z$ value to $2.6\mu \mathrm{m}$ from $0.77\mu \mathrm{m}$ [Figs. 2(k) and 2(l)]. Furthermore, at $z=0.6\mathrm{mm}$, the peak intensity dropped to 7% $I_0$ and the $\mathrm{FWHM}x$ and $\mathrm{FWHM}z$ values degraded to 0.37 and $3.9\mu \mathrm{m}$ [Fig. 2(m)], correspondingly. Besides, we confirmed that the spatial resolution and the signal, which degraded as a result of spherical aberrations, improved after correction with the continuous phase distribution placed at the front of the objective lens. [Figs. 2(e), 2(f), 2(i), and 2(j)]. At $z=0.3\mathrm{mm}$, the peak intensity recovered 93% $I_0$ [Fig. 2(l)]. On the other hand, we numerically calculated that the peak intensity of the PSF corrected by the 48-level quantized phase distribution was 96% of that corrected by the continuous phase distribution. Therefore, it was assumed that 89% $I_0$ was recovered after the correction with the quantized phase distribution at the front of the objective lens. Similarly, at $z=0.6\mathrm{mm}$, the peak intensity recovered by 85% compared to $I_0$ after the correction by the continuous phase distribution [Fig. 2(m)]. Furthermore, it was numerically calculated that the peak intensity of the PSF corrected by the quantized phase distribution dropped to 90% of that corrected by the continuous phase distribution. Therefore, it was assumed that the peak intensity recovered 77% $I_0$ after the correction by the quantized phase distribution at the front of the objective lens. On the other hand, at $z=0.3$ and 0.6 mm, the $\mathrm{FWHM}z$ values improved to 0.79 and $0.81\mu \mathrm{m}$, respectively, and the $\mathrm{FWHM}x$ values improved to $0.22\mu \mathrm{m}$ [Figs. 2(l) and 2(m)]. The FWHM values were almost equal to those at $z=0\mathrm{mm}$, after the correction by the continuous phase distribution.

Fig. 2

Numerically calculated point spread functions (PSFs) at a 488-nm wavelength. (a) Lateral and (b) axial images at $z=0\mathrm{mm}$. Lateral and axial images (c, d) before and (e, f) after correction at $z=0.3\mathrm{mm}$. Lateral and axial images (g, h) before and (i, j) after correction at $z=0.6\mathrm{mm}$. (k-m) Intensity profiles across the center of the intensity distribution along the $x$- and the $z$-axes at each depth $z$. The scale bars denote $0.5\mu \mathrm{m}$.

3.2.

Imaging of the Beads with Single-Photon Excitation Microscopy

Next, the correction effect of the device was experimentally evaluated by single-photon excitation laser scanning microscopy. In this experiment, fluorescent beads embedded in agarose gel were observed using 488-nm-wavelength laser light to evaluate the fluorescence signal and the FWHM of the image. The correction collar of the objective lens was adjusted beneath the cover glass’s surface ($z=0\mathrm{mm}$) to obtain the best optical performance. At a depth of 0.3 mm from the cover glass surface ($z=0.3\mathrm{mm}$), the peak intensity of the image dropped to 47% $I_0$, and the $\mathrm{FWHM}x$ value degraded to 0.38 from $0.27\mu \mathrm{m}$. The $\mathrm{FWHM}z$ value degraded to $2.8\mu \mathrm{m}$ from $1.6\mu \mathrm{m}$ [Figs. 3(a)–3(d), 3(k), and 3(l)]. Furthermore, at $z=0.6\mathrm{mm}$, the peak intensity dropped to 30% $I_0$. The $\mathrm{FWHM}x$ value degraded to $0.42\mu \mathrm{m}$, and the $\mathrm{FWHM}z$ value degraded to $5.0\mu \mathrm{m}$ [Figs. 3(g), 3(h), and 3(m)]. On the other hand, at $z=0.3\mathrm{mm}$, the peak intensity recovered by 116% compared to $I_0$, and the $\mathrm{FWHM}x$ value improved to $0.28\mu \mathrm{m}$ after correction with the device. Furthermore, the device improved the $\mathrm{FWHM}z$ value to $1.2\mu \mathrm{m}$ [Figs. 3(e), 3(f), and 3(l)]. Similarly, at $z=0.6\mathrm{mm}$, the device recovered 78% $I_0$, and improved the $\mathrm{FWHM}x$ value to $0.28\mu \mathrm{m}$. The $\mathrm{FWHM}z$ value improved to $1.5\mu \mathrm{m}$ after correction by the device [Figs. 3(i), 3(j), and 3(m)]. Namely, at $z=0.3$ and 0.6 mm, the peak intensities of the corrected images were approximately 2.5 and 2.6 times higher than those of the aberrated images, respectively. Furthermore, the FWHM values were similar to those at $z=0\mathrm{mm}$ after correction with the device.

Fig. 3

Images of fluorescent beads in single-photon excitation microscopy illuminated with 488-nm-wavelength laser light. The correction collar of the objective lens was adjusted to obtain the best performance beneath the cover glass surface ($z=0\mathrm{mm}$). (a) Lateral and (b) axial images at $z=0\mathrm{mm}$. Lateral and axial images (c, d) before and (e, f) after correction at $z=0.3\mathrm{mm}$. Lateral and axial images (g, h) before and (i, j) after correction at $z=0.6\mathrm{mm}$. (k-m) Intensity profiles across the center of the intensity distribution along the $x$- and the $z$-axes at each depth $z$. The data points in (k-m) are the mean values of several beads. The scale bars represent $0.5\mu \mathrm{m}$.

Next, the correction collar of the objective lens was adjusted to obtain the best performance at $z=0.7\mathrm{mm}$. At depths smaller than $z=0.7\mathrm{mm}$, the spatial resolution degraded and the fluorescence signal decreased [Figs. 4(a)–4(d)]. At $z=0.1\mathrm{mm}$, the peak intensity dropped to 58% of the peak intensity at $z=0.7\mathrm{mm}$ ($I_{0.7}$). Also, the $\mathrm{FWHM}x$ and $\mathrm{FWHM}z$ values degraded to $0.36\mu \mathrm{m}$ from $0.29\mu \mathrm{m}$ and to $3.3\mu \mathrm{m}$ from $1.8\mu \mathrm{m}$, respectively [Figs. 4(g) and 4(h)]. However, after correction with the device, the peak intensity at $z=0.1\mathrm{mm}$ recovered 94% $I_{0.7}$. Furthermore, the $\mathrm{FWHM}x$ value improved to $0.30\mu \mathrm{m}$, and the $\mathrm{FWHM}z$ value improved to $1.5\mu \mathrm{m}$ [Figs. 4(e), 4(f), and 4(h)]. The correction with the device thus increased the peak intensity by approximately 1.6 times compared to that of the aberrated image. Additionally, the FWHM values improved to the values almost equal to those at $z=0.7\mathrm{mm}$. From these results, it is evident that the device improved the spatial resolution and the fluorescence signal, and restored the cross-sectioning ability within a depth of $\pm 0.6\mathrm{mm}$.

Fig. 4

Images of fluorescent beads in single-photon excitation microscopy illuminated with 488-nm-wavelength laser light. The correction collar of the objective lens was adjusted to obtain the best performance at $z=0.7\mathrm{mm}$. (a) Lateral and (b) axial images at $z=0.7\mathrm{mm}$. Lateral and axial images (c, d) before and (e, f) after correction at $z=0.1\mathrm{mm}$. (g, h) Intensity profiles across the center of the intensity distribution along the $x$- and the $z$-axes at each depth $z$. The data points in (g, h) are the mean values of several beads. The scale bars denote $0.5\mu \mathrm{m}$.

3.3.

Imaging of the Beads with Two-Photon Excitation Microscopy

In preparation for the implementation of the device to observe a biospecimen using TPLSM, we introduced the device to a TPLSM illuminated by 950-nm-wavelength laser light and observed the fluorescent beads in order to confirm the correction effect of the device. The correction collar of the objective lens was adjusted to obtain the best performance at $z=0\mathrm{mm}$. As with single-photon excitation microscopy, the spatial resolution and the fluorescence signals of the images degraded in proportion to the depth where the beads were located [Figs. 5(a)–5(d), 5(g), and 5(h)]. At $z=0.3\mathrm{mm}$, the peak intensity dropped to 67% $I_0$. Also, the $\mathrm{FWHM}x$ and $\mathrm{FWHM}z$ values degraded to $0.50\mu \mathrm{m}$ from $0.44\mu \mathrm{m}$, and to $3.9\mu \mathrm{m}$ from $2.3\mu \mathrm{m}$, respectively [Figs. 5(k) and 5(l)]. Furthermore, at $z=0.6\mathrm{mm}$, the peak intensity dropped to 45% $I_0$. The $\mathrm{FWHM}x$ value degraded to $0.55\mu \mathrm{m}$, and the $\mathrm{FWHM}z$ value degraded to $4.5\mu \mathrm{m}$ [Fig. 5(m)]. On the other hand, at $z=0.3$ and 0.6 mm, the device recovered 115% $I_0$ and 98% $I_0$, respectively. The $\mathrm{FWHM}z$ values at $z=0.3$ and 0.6 mm improved to 1.9 and $2.1\mu \mathrm{m}$, correspondingly, and the $\mathrm{FWHM}x$ values to $0.44\mu \mathrm{m}$ [Figs. 5(e), 5(f), 5(i), 5(j), 5(l), and 5(m)]; namely, the correction by the device increased the peak intensity by approximately 1.7 and 2.2 times, respectively, compared to the aberrated conditions. The FWHM values improved to the value almost equal to those at $z=0\mathrm{mm}$. At $z=0.3$ and 0.6 mm, the phase amplitudes generated by the device were 1550 and 3300 nm, respectively. In this way, we succeeded in improving the lateral and axial resolutions. Additionally, the fluorescence signal in deep regions was improved like in the case of single-photon excitation microscopy.

Fig. 5

Images of fluorescent beads in two-photon excitation microscopy illuminated with 950-nm-wavelength laser light. The correction collar of the objective lens was adjusted to obtain the best performance beneath the cover glass surface ($z=0\mathrm{mm}$). (a) Lateral and (b) axial images at $z=0\mathrm{mm}$. Lateral and axial images (c, d) before and (e, f) after correction at $z=0.3\mathrm{mm}$. Lateral and axial images (g, h) before and (i, j) after correction at $z=0.6\mathrm{mm}$. (k-m) Intensity profiles across the center of the intensity distribution along the $x$- and the $z$-axes at each depth $z$. The data points in (k-m) are the mean values of several beads. The scale bars represent $0.5\mu \mathrm{m}$.

3.4.

Imaging of the Biospecimen with Two-Photon Excitation Microscopy

Finally, we implemented the device for the observation of the biospecimen described earlier using TPLSM. Figures 6 and 7 show the images of the TDE-treated fixed mouse brain slice. The observation regions were at the surface of the cover glass ($z=0\mathrm{mm}$) and at depths 0.19, 0.27, and 0.54 mm from the surface. The same areas were observed before and after correcting aberrations. As with the observation of the bead sample, we utilized dendritic spines about $1\mu \mathrm{m}$ in size as a ruler to estimate the spatial resolution in the fixed brain slice.

Fig. 6

Images of the biospecimen in two-photon excitation microscopy illuminated with 950-nm-wavelength laser light. Lateral images (a) before and (b) after correction by the device at $z=0.19\mathrm{mm}$, when the correction collar of the objective lens was adjusted to obtain the best performance beneath the cover glass surface ($z=0\mathrm{mm}$). (c) Lateral image at $z=0.19\mathrm{mm}$, when the correction collar of the objective lens was adjusted to obtain the best performance at $z=0.19\mathrm{mm}$. Lateral images (d) before and (e) after correction by the device at $z=0.27\mathrm{mm}$, when the correction collar of the objective lens was adjusted to obtain the best performance at $z=0\mathrm{mm}$. (f) Lateral image at $z=0.27\mathrm{mm}$, when the correction collar of the objective lens was adjusted to obtain the best performance at $z=0.27\mathrm{mm}$. The images in (g-r) were obtained in the area indicated by a rectangle in (a-f). (g) Lateral and (h) axial images before and after correction by (i, j) the device and (k, l) the correction collar at $z=0.19\mathrm{mm}$. Lateral and axial images (m, n) before and after correction by (o, p) the device and (q, r) the correction collar at $z=0.27\mathrm{mm}$. (s, t) Intensity profiles across the center of the intensity distribution along the $x$- and the $z$-axes at each depth $z$. The scale bars denote $5\mu \mathrm{m}$ in (a) and $0.5\mu \mathrm{m}$ in (g) and (h).

Fig. 7

Images of the biospecimen in two-photon excitation microscopy illuminated with 950-nm-wavelength laser light. The correction collar of the objective lens was adjusted to obtain the best performance at $z=0.27\mathrm{mm}$. Lateral images (a) before and (b) after correction by the device at $z=0\mathrm{mm}$. Lateral images (c) before and (d) after correction by the device at $z=0.54\mathrm{mm}$. The images (e-l) were obtained at the area indicated by the rectangle in (a-d). Lateral and axial images (e, f) before and (g, h) after correction by the device at $z=0\mathrm{mm}$. Lateral and axial images (i, j) before and (k, l) after correction by the device at $z=0.54\mathrm{mm}$. (m, n) Intensity profiles across the center of the intensity distribution along the $x$- and the $z$-axes at each depth $z$. The scale bars denote $5\mu \mathrm{m}$ in (a) and $0.5\mu \mathrm{m}$ in (i) and (j).

Figure 6 shows the images acquired when the correction collar of the objective lens was adjusted to obtain the best performance at $z=0\mathrm{mm}$. From the lateral wide-field images at $z=0.19$ and 0.27 mm, the fluorescence signals improved after correction by the device, as well as with the correction collar [Figs. 6(a)–6(f)]. Furthermore, from the images of the dendritic spines, the device and the correction collar clearly improved the spatial resolution and the fluorescence signal [Figs. 6(g)–6(r)]. Specifically, at $z=0.19\mathrm{mm}$ and a phase amplitude of 2300 nm, the device improved the $\mathrm{FWHM}x$ value to $0.48\mu \mathrm{m}$ from $0.65\mu \mathrm{m}$ and the $\mathrm{FWHM}z$ value to $1.4\mu \mathrm{m}$ from $3.3\mu \mathrm{m}$. Furthermore, the peak intensity of the image corrected by the device was approximately 1.8 times higher than that of the aberrated image [Fig. 6(s)]. On the other hand, the correction collar improved the $\mathrm{FWHM}x$ value to $0.53\mu \mathrm{m}$ and the $\mathrm{FWHM}z$ value to $1.8\mu \mathrm{m}$. The peak intensity of the image corrected by the correction collar was approximately 2.1 times higher than that of the aberrated image [Fig. 6(s)]. At $z=0.27\mathrm{mm}$, the device and the correction collar improved the $\mathrm{FWHM}x$ value to $0.65\mu \mathrm{m}$ from $0.74\mu \mathrm{m}$, they improved the $\mathrm{FWHM}z$ value to 1.5 and $1.7\mu \mathrm{m}$ from $4.0\mu \mathrm{m}$, respectively, and increased the peak intensity by approximately 2.4 times compared to that of the aberrated condition [Fig. 6(t)]. At $z=0.27\mathrm{mm}$, the phase amplitude of the device was 2900 nm. In the experiment with the bead sample, the correction by the device with phase amplitude 3300 nm increased the peak intensity 2.2 times compared to that of the aberrated condition (see Sec. 3.3). Accordingly, for the biospecimen, the device showed a similar improvement of the fluorescence signal to that observed for the bead sample. From this result, we confirmed that the device improved the lateral and the axial resolution; additionally, the fluorescence signal degraded in deep regions of the biospecimen, as with the bead sample. Furthermore, the device showed a correction effect similar to that of the correction collar.

Finally, in order to observe a wider range of depth, we adjusted the correction collar to obtain the best performance at a deep region of the biospecimen. In Sec. 3.2, we showed that the device improved the image quality at both shallower and deeper regions than the position in which the best optical performance was obtained. On the other hand, the device improved the image quality at $z=0–0.27\mathrm{mm}$ when the correction collar was adjusted to obtain the best performance at $z=0\mathrm{mm}$ (Fig. 6). Meanwhile, the maximum depth at which the depth-induced spherical aberration was correctable by the correction collar was 0.27 mm. By utilizing this property of the device and the objective lens, the device was expected to improve the image quality at $z=0–0.54\mathrm{mm}$ when the correction collar was adjusted to obtain the best performance at $z=0.27\mathrm{mm}$. Thus, we observed the biospecimen at depths of 0–0.54 mm. As in Fig. 6, at $z=0$ and 0.54 mm, the fluorescence signal of the wide-field lateral images improved after correction by the device [Figs. 7(a)–7(d)]. From the images of the dendritic spines, at $z=0\mathrm{mm}$, the $\mathrm{FWHM}x$ and $\mathrm{FWHM}z$ values improved to $0.57\mu \mathrm{m}$ from $0.66\mu \mathrm{m}$, and to $2.1\mu \mathrm{m}$ from $3.2\mu \mathrm{m}$, respectively. Additionally, the peak intensity of the image corrected by the device was approximately 1.5 times higher than that of the aberrated image [Figs. 7(e)–7(h), 7(m)]. Furthermore, at $z=0.54\mathrm{mm}$, the $\mathrm{FWHM}x$ value improved to $0.60\mu \mathrm{m}$ from $0.70\mu \mathrm{m}$ and the peak intensity of the corrected image was approximately 1.9 times higher than that of the aberrated image. At $z=0.54\mathrm{mm}$, before correction, the intensity profile along the $z$-axis extended outside the observation range and the $\mathrm{FWHM}z$ value was estimated from the fitting curve. Therefore, the $\mathrm{FWHM}z$ value improved to $2.0\mu \mathrm{m}$ from $4.5\mu \mathrm{m}$ after correction by the device [Figs. 7(i)–7(l), 7(n)]. Thus, by using the device, we succeeded in the acquisition of a high-contrast and high-resolution image of the biospecimen in a wide range of depths, including a depth twice as large as that where depth-induced spherical aberration was correctable by the correction collar.