3.1.

Confocal Fluorescence Imaging over a Microtome Knife

In the fMOST system, the sliced brain sample strips are imaged over the diamond knife. For a brain sample with fluorescently labeled neurons distributed across the whole brain, the brain samples under the diamond knife will also be excited (the diamond knife is not opaque to the excitation and emission light) and emit fluorescence. The fluorescence will appear as background noise in the sample strip image, and this noise should be suppressed to enhance the image contrast. For the imaging while sectioning method employed in the proposed system, the sample strips being imaged are sliced and separated from the sample under the knife, and the excitation light cone is elevated relative to the base of the sample, as illustrated in Fig. 2. This mechanical separation provides some degree of background suppression. A confocal scheme is employed to provide additional suppression of the fluorescence background. That is, we can tune the fluorescence background suppression by adjusting the optical sectioning capability (by changing the confocal slit width) or change the elevation distance of the excitation light cone by changing the imaging site over the diamond knife.

Fig. 2

Fluorescence excitation and emission in confocal fMOST system. (a) Light cone excites not only the sample strip over the knife but also the sample below, introducing background noise. There is a distance $D$ between the sample strip and the bulk sample below. Moving the light cone along the knife increases the distance $D$, and the excited fluorescence background will decrease. (b) The signal-to-noise ratio (SNR) increases with $D$, as illustrated. (c) Using a confocal slit, the detection volume will be limited, and the fluorescence background will be reduced. (d) The relationship between SNR and slit width is illustrated.

To further evaluate the effect of confocal detection and spatial separation (light cone elevation distance) on fluorescence background suppression, we constructed a simplified ideal model, as shown in Fig. 3, and considered an infinitely small pinhole for confocal detection analysis.

Fig. 3

Schematic diagram of a simplified imaging model. A $2\text{-}\mu \mathrm{m}$-thick slice is placed over the bulk sample with an elevation distance $d$. The objective lens is simplified as a circular aperture of radius $a$ with a focal length of $z_f$. $P_f$ is the focal point, $P_d$ is a point near the focal point. By summing up the contribution from the points in the bulk sample, background intensity contribution from the bulk sample at different elevation distances can be analyzed. $\delta z=0$ indicates the focal plane of the objective lens. The value of intensity point-spread function (PSF) at $\delta z>30\mu \mathrm{m}$ is very small and is neglected in the model.

We use Born and Wolf analysis of the circular aperture diffraction problem^19,20 (modeling the objective lens). As shown in Fig. 3, the objective lens is simplified to a circular aperture of radius $a$ with focal length $z_f$. The field at $P_d(x_d,y_d,z_d)$ near the focal point $P_f(\mathrm{0,0},z_f)$ of a spherical wave can be regarded as the expression of the excitation point-spread function (PSF). In the objective lens, we use the approximation $\mathrm{NA}\approx a/z_f$. The complex amplitude at $P_d$ is²¹

The PSF at point $P_d$ can be written as $h(P_d)=U(P_d)$. With the same objective lens used for both excitation illumination and emission detection, we can assume the illumination and detection PSF follow the same distribution function, ignoring the difference in wavelength. The imaging intensity PSF can be written as²¹

As shown in Fig. 3, a $2\text{-}\mu \mathrm{m}$-thick slice is placed over the bulk sample with an elevation distance $d$. The PSF reflects the contribution of each point in the sample space to the detected intensity, assuming a uniform plane wave illumination parallel to the optic axis. The intensity PSF makes a negligible contribution when $r>10\mu \mathrm{m}$ or $\delta z>30\mu \mathrm{m}$. Because the PSF is radially symmetric near the focal point, we consider only the $xoz$ plane, and calculate the intensity PSF inside the regions defined by $-10\mu \mathrm{m}\le x\le 10\mu \mathrm{m}$ and $0\mu \mathrm{m}\le \delta z\le 30\mu \mathrm{m}$. The PSF is represented as a matrix ${(a_{ij})}_{301\times 201}$, $i=\mathrm{1,2},\dots ,301$; $j=\mathrm{1,2},\dots ,201$. The elements of the matrix represent the intensity contribution from a $0.1\mu \mathrm{m}\times 0.1\mu \mathrm{m}$ area of the detected signal.

The signal from a slice is calculated as $\sum _{1\le j\le 201}^{i=1}{a}_{ij}+2\sum _{1\le j\le 201}^{2\le i\le 10}{a}_{ij}$, and the fluorescence background from the bulk sample can be calculated as $\sum _{1\le j\le 201}^{n+10<i\le 301}{a}_{ij}$, $n=d/0.1$. The SNR is defined as

We first consider the nonconfocal condition, in which the confocal pinhole is not used. Figures 4(a)–4(c) show the background intensity from the bulk sample at various elevation distances ($d=0\mu \mathrm{m}$, $5\mu \mathrm{m}$, and $15\mu \mathrm{m}$) summed along the $z$-axis. Figure 4(d) shows the signal intensity from the slice, summed along the $z$-axis. Figure 4(e) shows that SNR increases with elevation distance. The increase in SNR at different elevation distances relative to the SNR at $d=0$ has a hint of the background suppression effect of slice separation by mechanical sectioning. The SNR is about 1.6 at $d=0$, and increases by a factor of 2.4, 4.3, and 8.3 at $d=5\mu \mathrm{m}$, $10\mu \mathrm{m}$, and $15\mu \mathrm{m}$, respectively. In practice, the short pixel integration time at high imaging speeds requires high excitation power to enhance the signal intensity. This will generate a strong fluorescence background from the bulk sample beneath the diamond knife. The increase in SNR provided by physical separation alone does not provide sufficient suppression of the strong fluorescence background.

Fig. 4

Results for nonconfocal detection condition. (a, b, c) The background intensity contribution from the bulk sample (contribution along $x$-axis) at $0\mu \mathrm{m}$, $5\mu \mathrm{m}$, and $15\mu \mathrm{m}$ elevation distances. (d) The signal intensity from the sample slice (along $x$-axis). (e) SNR at different elevation distances $d$.

Figure 5 shows the results for the confocal condition when the detection pinhole is used. Figures 5(a)–5(c) show the background intensity from the bulk sample at increased elevation distances ($d=0\mu \mathrm{m}$, $5\mu \mathrm{m}$, and $15\mu \mathrm{m}$) summed along the $z$-axis. Figure 5(d) shows the signal intensity from the slice summed along the $z$-axis. In Fig. 5(e), the SNR is about 79 at zero elevation. The SNR increases 52, 224, and 728 fold at $5\mu \mathrm{m}$, $10\mu \mathrm{m}$, and $15\mu \mathrm{m}$ elevation distances, respectively. The SNR increase resulting from the increase in elevation distance is much more obvious in the confocal condition than in the nonconfocal condition. Although this is a much simplified and idealized simulation, it suggests that, compared with normal confocal imaging, physical separation of the slice from the bulk sample below can greatly improve the suppression of fluorescence background, even with a small separation distance. Compared with the above results for the nonconfocal condition, the improvement in SNR with confocal detection is quite significant.

Fig. 5

Result for confocal detection condition. (a, b, c) The background intensity contribution from the bulk sample (contribution along $x$-axis) at $0\mu \mathrm{m}$, $5\mu \mathrm{m}$, and $15\mu \mathrm{m}$ elevation distances. (d) The signal intensity from the sample slice (along $x$-axis). (e) SNR at different elevation distances $d$.

We illustrate the effect of combining confocal detection with physical separation on background suppression in Fig. 6. Resin-embedded mouse brain samples were used. Figures 6(a)–6(c) show the results from a thy1-GFP-M mouse brain sample. Figures 6(d)–6(f) show the result from a thy1-YFP-H mouse brain sample. For the experiments, a sample strip of thickness $2\mu \mathrm{m}$ containing a few fluorescent neurites was first sliced and collected. A confocal image of the sample strip was acquired using a Zeiss LSM 780 confocal microscope, with a $20\times /1.0$ objective lens and $50\mu \mathrm{m}$ pinhole diameter. Using the same imaging parameters as for the sample strip, confocal images of the remaining bulk sample block were then acquired with the objective focal plane set to $2\mu \mathrm{m}$ and $10\mu \mathrm{m}$ above the block surface, respectively. Figure 6 shows an overlay of these two images with the sample image. This can be regarded as a simulation of imaging with only the confocal scheme and imaging combining confocal detection with a small separation between the sample strip and the bulk sample. As seen in Figs. 6(a) and 6(d), residual fluorescence background from the cell body of neurons at the block surface can be visually identified. Intensity profile along the specified line [blue line in Figs. 6(a) and 6(d), and red line in Figs. 6(b) and 6(e)] shows that residual signal intensity from the cell body, which is located at the surface of the sample block, is close to or even larger than the signal from the neurites in the sample strip. Note that the fluorescence of the cell body can be several times brighter than that of the neurites. This accounts for the above fact that, even located at the defocused plane of a confocal microscope, the residual signal intensity from the cell body is still close to that from the neurites imaged at the focal plane. Figures 6(b) and 6(e) show that a $10\mu \mathrm{m}$ elevation of the focal plane (the light cone) largely suppresses the background. An image stack with the objective focal plane placed at the surface and above was also acquired and the residual signal from the cell body [indicated by the white arrow in Figs. 6(a) and 6(d)] at different elevation distances $d$ are shown in the inset of Figs. 6(c) and 6(f). The normalized signal intensity drops from around 0.5 [0.44 and 0.57 respectively, in Figs. 6(c) and 6(f)] at $d=2\mu \mathrm{m}$ to around 0.1 at $d=10\mu \mathrm{m}$ [0.08 and 0.12 respectively, in Figs. 6(c) and 6(f)]. This demonstrates the necessity of combining confocal detection and physical separation for this type of neural wiring imaging application, where signal intensity from the cell body and that from the neurite can span an order of magnitude.

Fig. 6

Images given by overlaying confocal image of a sliced $2\text{-}\mu \mathrm{m}$-thick sample with a few fluorescent neurites inside and confocal image of the remaining bulk sample block with focal plane of the objective at a distance of $2\mu \mathrm{m}$ [(a, d)] and $10\mu \mathrm{m}$ [(b, e)] above block surface. Scale bar: $20\mu \mathrm{m}$. Upper row shows the results from a thy1-GFP-M mouse brain and lower row shows the results from a thy1-YFP-H mouse brain. (c, f) The intensity profile along the specified lines in the images. (a, d) Black arrows indicate residual signal from cell bodies marked by the white arrows. (c, f) Inset plots show the residual signal from these two cell bodies at different elevation distances $d$ of the objective focal plane.

Confocal detection can also help to recover the signal from the small tears at the edge of adjacent strips that sometimes occur in the sectioning process.²² The signal will otherwise be buried in the fluorescence background.

3.2.

Fast and Stable Scanning

Fast and stable imaging is important for the acquisition of a complete whole-brain image dataset at submicron voxel size. For the point scan confocal imaging method, this requires a fast and stable scanner. As the time required to image a mouse brain could be more than 1 week with the point scan method, an inertia-free scanner based on the acousto-optical principle was chosen.^18,23

The imaging of each sample strip requires the coordinated operation of the AOD scanner and the sample stage. AOD only produces a 1-D scan. For two-dimensional imaging, a scan along another dimension is given by sliding the sample strip over the knife (i.e., by moving the sample stage along the $x$-axis) (see Fig. 1).

To perform fast scanning, the AOD scanner works in a high-frequency chirp mode.^18,19 However, in this mode, the laser beam, in addition to being deflected, will also diverge or converge in only one direction, which will in turn cause astigmatism. Working in high-frequency chirp mode, the AOD is actually equivalent to a cylindrical acousto-optical lens (AO lens) with a focal length equal to $f=v^2/\lambda \alpha$,^24–26 where $v$ is the propagation speed of an acoustic wave inside the AOD crystal, $\lambda$ is the wavelength of the excitation light, and $\alpha$ is the rate of change of the acoustic frequency. The cause of astigmatism and its influence on imaging resolution are shown in Fig. 7. We can correct this effect by introducing a cylindrical lens (CL), which is placed after the AOD at a distance equal to the focal length difference of the AO lens and the introduced CL. The AOD used in our implementation has a frequency bandwidth of 50 MHz, and the scan time per line was set to $150\mu \mathrm{s}$. The acoustic wave propagation speed is $650\mathrm{m}/\mathrm{s}$, and the calculated focal length of the AO lens is longer than 2 m. We made a custom CL with a focal length (absolute value) 30 mm shorter than that of the AO lens at 473 nm. With such long focal lengths of both the AO lens and the CL, we found that the distance between the two became less important, and can be fixed even at different excitation wavelengths ($473/488\mathrm{nm}$ for GFP or 515 nm for YFP), although the resolution can be significantly improved, as shown in Fig. 8(b).

Fig. 7

Astigmatism caused by AOD working in the high-frequency chirp scan mode. In this mode, AOD is equivalent to a cylindrical lens (AO lens). Astigmatism occurs when the beam is focused by the objective lens.

Fig. 8

A cylindrical lens can be introduced to correct the AOD astigmatism problem. (a) The CL was placed after the AOD. The distance between the two is equal to their focal length difference. (b) Spatial resolution and fluorescence intensity increased ($\sim 2$ times) after the CL was introduced. Data from images of 170 nm fluorescent beads.

One disadvantage of AOD is the limited scan angle. We used a scan lens (L1 in Fig. 1) with a long focal length (400 mm), so that the scan angle at the back aperture of the objective lens will be 2.2 times that of the AOD. In this case, the beam diameter will also shrink. To make full use of the objective numerical aperture, we chose an AOD with a large aperture (15 mm), and maintained a 6.8 mm beam diameter at the back aperture of the objective lens. With the above configuration, a $375\mu \mathrm{m}$ scan length (488 nm excitation, $40\times$ objective lens) can be achieved with submicron spatial resolution [Fig. 7(b), red line].