Wide-field optical sectioning for live-tissue imaging by plane-projection multiphoton microscopy

Article (PDF Available)inJournal of Biomedical Optics 16(11):116009 · November 2011with21 Reads
DOI: 10.1117/1.3647570 · Source: PubMed
Abstract
Optical sectioning provides three-dimensional (3D) information in biological tissues. However, most imaging techniques implemented with optical sectioning are either slow or deleterious to live tissues. Here, we present a simple design for wide-field multiphoton microscopy, which provides optical sectioning at a reasonable frame rate and with a biocompatible laser dosage. The underlying mechanism of optical sectioning is diffuser-based temporal focusing. Axial resolution comparable to confocal microscopy is theoretically derived and experimentally demonstrated. To achieve a reasonable frame rate without increasing the laser power, a low-repetition-rate ultrafast laser amplifier was used in our setup. A frame rate comparable to that of epifluorescence microscopy was demonstrated in the 3D imaging of fluorescent protein expressed in live epithelial cell clusters. In this report, our design displays the potential to be widely used for video-rate live-tissue and embryo imaging with axial resolution comparable to laser scanning microscopy.
Wide-field optical sectioning for live-tissue
imaging by plane-projection multiphoton
microscopy
Jiun-Yann Yu
Chun-Hung Kuo
Daniel B. Holland
Yenyu Chen
Mingxing Ouyang
Geoffrey A. Blake
Ruben Zadoyan
Chin-Lin Guo
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Journal of Biomedical Optics 16(11), 116009 (November 2011)
Wide-field optical sectioning for live-tissue imaging
by plane-projection multiphoton microscopy
Jiun-Yann Yu,
a
Chun-Hung Kuo,
b
Daniel B. Holland,
c
Yenyu Chen,
d
Mingxing Ouyang,
a
Geoffrey A. Blake,
c,e
Ruben Zadoyan,
b
and Chin-Lin Guo
a
a
California Institute of Technology, Bioengineering, 1200 E. California Boulevard, MC 138-78, Pasadena,
California 91125
b
Newport Corporation, Technology and Applications Center, 1791 Deere Avenue, Irvine, California 92606
c
California Institute of Technology, Division of Chemistry and Chemical Engineering, 1200 E. California Boulevard,
MC 139-74, Pasadena, California 91125
d
Stanford University, Center for Cardiovascular Technology, 300 Pasteur Drive, Palo Alto, California 94305
e
California Insitute of Technology, Division of Geological and Planetary Sciences, 1200 E. California Boulevard, MC
150-21, Pasadena, California 91125
Abstract. Optical sectioning provides three-dimensional (3D) information in biological tissues. However, most
imaging techniques implemented with optical sectioning are either slow or deleterious to live tissues. Here, we
present a simple design for wide-field multiphoton microscopy, which provides optical sectioning at a reasonable
frame rate and with a biocompatible laser dosage. The underlying mechanism of optical sectioning is diffuser-based
temporal focusing. Axial resolution comparable to confocal microscopy is theoretically derived and experimentally
demonstrated. To achieve a reasonable frame rate without increasing the laser power, a low-repetition-rate ultrafast
laser amplifier was used in our setup. A frame rate comparable to that of epifluorescence microscopy was
demonstrated in the 3D imaging of fluorescent protein expressed in live epithelial cell clusters. In this report, our
design displays the potential to be widely used for video-rate live-tissue and embryo imaging with axial resolution
comparable to laser scanning microscopy.
C
2011 Society of Photo-Optical Instrumentation Engineers (SPIE). [DOI: 10.1117/1.3647570]
Keywords: microscopy; multiphoton processes; diffusers; imaging systems.
Paper 11137RR received Mar. 22, 2011; revised manuscript received Sep. 3, 2011; accepted for publication Sep. 19, 2011; published
online Nov. 3, 2011.
1 Introduction
Optical sectioning and high-acquisition-rate imaging techniques
have been established for decades, but very few techniques have
been proposed to provide optical sectioning at a reasonable
frame rate and with a biocompatible laser dosage for live-tissue
imaging. Most optical-sectioning techniques rely on a scanned
optical probe point.
1
However, to obtain images at a reasonable
frame rate, the excitation intensity of laser scanning microscopy
often has to be significantly stronger than wide-field microscopy;
this is due to the extremely short dwell time per pixel in laser
scanning microscopy. Consequently, for confocal microscopy,
significant phototoxicity can occur in scanned live organisms if a
video-rate time-lapse microscopy is required.
1
Such phototoxi-
city can be greatly reduced by using multiphoton microscopy,
2, 3
but a tradeoff is the thermal mechanical damage to live tissue
through the single-photon absorption of the near-infrared
excitation.
4
One potential way to resolve photo/thermal-damage in scan-
ning microscopy without the loss of acquisition speed is
to implement the capability of optical sectioning in wide-
field microscopy. Several methods have been proposed to
fulfill this goal.
59
These methods can be categorized into
two main regimes: single-photon excitation and multiphoton
excitation.
Address all correspondence to: Jiun-Yann Yu, California Institute of Technology,
1200 E California Blvd, MC 138-78, Pasadena, CA 91125; Tel: 626-395-5992;
E-mail: jyyu@caltech.edu.
In wide-field single-photon excitation, the most well-known
methods include light-sheet illumination microscopy and
structured light microscopy.
5, 6
Both methods are technically
elaborate, introducing complexity into the optical system due to
the requirement of additional mechanical parts that synchronize
with axial scanning components. Moreover, light-sheet mi-
croscopy obtains optical sectioning by illuminating the sample
laterally (i.e., perpendicular to the optical axis); this introduces
significant mechanical complexity and makes the preparation
of samples difficult. Furthermore, structured light microscopy
excites the full sample volume and requires multiple exposures
for each single position—an extremely inefficient use of the
quantum yield of the fluorophores.
6
For wide-field multiphoton excitation, a key element to main-
tain the capability of optical sectioning is to create significant
time delays among positions excited in the illumination field,
7, 8
as we discuss in Sec. 2. Consequently, the complexity of the
systems is further increased by the need to engineer sufficient
time delays, and the number of pixels has been very limited
(<100).
7, 8
The concept of time delays in wide-field multiphoton
excitation was later generalized into an optical design referred
to as temporal focusing.
9
In a temporal focusing setup, the
pulse width of the excitation light varies in time as the pulse
propagates a long the optical axis, and is shortest at the image
plane (Fig. 1 inset); the higher peak intensity associated with
a shorter pulse width provides the optical sectioning effect.
1083-3668/2011/16(11)/116009/9/$25.00
C
2011 SPIE
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Yu et al.: Wide-field optical sectioning by plane-projection multiphoton microscopy
Fig. 1 The setup of a plane-projection multiphoton microscope. The
ultrafast infrared laser beam (IR, red) from the left is scattered by a
diffuser (DF). The image of the surface of DF is then projected to the
sample to excite a thin plane (green). The zoomed-in inset shows that
light arriving at an out-ot-focus point q
is mainly from points within
a cone angle θ on the image plane, and that the different arrival times
elongate the effective pulse width at q
.L
D
: diffuser lens of focal length
f
D
. B: beamsplitter. L
O
: microscope objective of focal length f
O
.S:
sample; L
I
: imaging lens. IP: image plane.
Temporal focusing for microscopy was first experimentally
demonstrated by Oron et al.
9
In their setup, the laser pulse
is directed to a blazed grating (which serves as a scatterer)
in an oblique incidence orientation. This creates a line-scan
mechanism, which leads to the time delay across the image
plane, thereby generating optical sectioning.
9
There exist two
issues, however, in this technique. First and primarily, a unique
blazed diffraction grating must be fabricated for each specific
wavelength window. Exciting multiple fluorophores, as is
required for most biomedical studies, would require multiple
gratings and hence increases the complexity of the system.
Second, the signal level obtained by the previous temporal
focusing setup
9
is weaker by several orders of magnitude
than that which can be achieved by conventional scanning
microscopy. The low signal level results from the dilution of
the excitation intensity in a wide-field setup. Consequently, the
image acquisition rate is significantly reduced. For example,
in Oron et al.’s original report,
9
the frame rate is roughly
0.033 frames per second (fps) for cells stained with DAPI
(a fluorescent dye for chromosome staining), which is much
brighter than most biological fluorophores expressed in live
tissues.
Here, we present a simple approach to remove the limita-
tions associated with a single excitation wavelength and low
acquisition rates in temporal-focusing microscopes. One way
to overcome the limitation o f single wavelength excitation is
to use an optical diffuser rather than a blazed grating as the
scatterer, because the scattering pattern of an optical diffuser
is insensitive to the central wavelength of the excitation light.
A theoretical estimation by Oron et al., however, suggests that
using optical diffusers to create temporal focusing requires the
pulse width of the laser to be shorter than 10 fs, even with
high numerical-aperture (NA) objectives.
9
This would make
diffuser-based temporal focusing almost impractical, given the
current pulse width of most commercially available light sources
(100 fs). In their estimations, though, the optical diffuser was
considered as an ideally flat plane of points generating ultrafast
pulses simultaneously. In practice, a ground-glass diffuser has
a rough surface, which introduces random time delays among
the scattering points. In other words, the optical diffuser cre-
ates a plane of point sources having a distribution of time de-
lays with respect to each other, instead of zero time delay as
was previously modeled. By projecting these point sources onto
the sample plane of the microscope, temporal focusing, and
hence optical sectioning, can be achieved. Therefore, we refer
to our technique as “plane-projection multiphoton” (PPMP) mi-
croscopy. Through geometrical calculations, we found that using
an optical diffuser should enable optical sectioning comparable
to confocal microscopy, even with moderate NA objectives and
pulse widths up to 100 fs.
For the other issue of temporal focusing, the low acquisition
rate, one needs a way to increase the signal level. Recent studies
suggest that using ultrafast laser amplifiers with a repetition rate
of approximately a few hundred kilohertz might significantly
enhance the signal level, without increasing the average power.
9
To examine this possibility, we theoretically derived the depen-
dence of the multiphoton excitation upon the pulse repetition
rate while the average power remained constant. Our calcula-
tions suggest that for a field of view comparable to conventional
epifluorescence microscopy, a repetition rate of approximately
1 kHz or slower could yield sufficient signal at a reasonable
frame rate and with a biocompatible laser dosage for live-cell
imaging. We experimentally verified this prediction by using
a 1-kHz ultrafast amplifier to obtain the optical sectioning of
fluorescent protein expressed in live epithelial tissues at a frame
rate of 5 fps, a frame rate similar to that used in a conventional
epifluorescence microscope to obtain images on the same sam-
ple. Thus, by using optical diffusers having sufficient surface
roughness, and laser sources of sufficiently low repetition rate
and high pulse energy, we have demonstrated a simple system
design for obtaining 3D live-tissue images at an axial resolution
comparable to conventional confocal microscopy. Furthermore,
signal levels, and thus frame rates, are comparable with conven-
tional epifluorescence microscopy.
2 Theory
2.1 Efficiency of Temporal Focusing Through
an Optical Diffuser
In this section we reinvestigate the temporal focusing effect
through an optical diffuser. Specifically, we compute the vari-
ation of the pulse width along the optical axis based on geo-
metrical optics. An axial resolution comparable to conventional
confocal microscopy is derived.
Figure 1 depicts the schematic of a typical temporal focusing
setup.
9
An optical diffuser (DF) was used to transform the in-
coming ultrafast laser beam into a plane of point sources. These
point sources were then projected onto the image plane (IP) of
an infinity-corrected microscope, through the diffuser lens (L
D
)
and the objective (L
O
). The resulting fluorescence was imaged
onto a CCD camera through an imaging lens (L
I
).
As discussed by Oron et al., the elongation of pulse width at
an out-of-focus point q
at a distance z away from the IP can
be estimated through the maximal difference of pulse arrival
times from the point sources within a cone of angle θ from the
IP (Fig. 1 inset).
9
Here, θ can be determined by the divergence
angle of L
O
, θ NA/n (Fig. 1 inset).
9
To estimate the difference
of pulse arrival times resulting from the geometry of the setup,
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Yu et al.: Wide-field optical sectioning by plane-projection multiphoton microscopy
Fig. 2 Random time delay generated by the surface roughness of an
optical diffuser. (a) The time delay of the laser pulses results from a
discrepancy of height, h, on the diffuser surface. (b) A more general
case for (a). Here, h corresponds to the maximal surface height dis-
crepancy (i.e., the peak-to-valley difference) within an area of radius r
(the corresponding area on the IP is of radius rf
O
/f
D
). (c) The schematic
illustration of h under different scales of r. Upper: when r 0,
h 0. Lower: when r D, h D.
we first considered the case where the DF is approximated
as a flat plane of point sources that simultaneously generate
ultrafast pulses. Using the lens formula and paraxial approx-
imation, the elongation of the pulse width
t
G
at the point q
can be estimated as
t
G
(z)
( f
D
+ f
O
d) · NA
2
2 C
0
· n · f
2
O
· z
2
+ n
n
n
2
NA
2
C
0
·
n
2
NA
2
· z,
(1)
where C
0
is the speed of light in vacuum and n is the refractive
index of the sample medium. The first term on the right-hand
side arises from the difference of optical path lengths from the
DF to IP, and the second term results from the difference of
optical path lengths from IP to the point q
.
We next take into account the surface roughness of DF and
estimate how such roughness leads to a randomness of arrival
times. To proceed, we consider a simple case where the rough-
ness is represented by a step function with a height discrepancy
h on the diffuser surface [Fig. 2(a)]. The time delay of a pulse
p1 induced by h is simply h/C
Glass
h/C
Air
= (n
Glass
n
Air
) · (h/C
0
) 0.5 h / C
0
. Further, the time delay caused
by the roughness in a region A
r
of radius r on the diffuser sur-
face is projected into a region A
r
of radius r
= f
O
/ f
D
r on the
image plane [Fig. 2(b)]. The maximal time delay t
RD
within
A
r
may thus be approximated as
t
RD
= 0.5
h
C
0
, (2)
where h is the maximal surface height discrepancy within A
r
.
In general, the roughness of a ground-glass diffuser is generated
by grinding a fl at surface of glass with particles of size less than
a certain length D. Thus, we expect h 0whenr 0, and
h D if r D, as shown in Fig. 2(c). To take into account
these asymptotic estimations, we used a simple approximation
here: h α · 2r if α · 2r < D and h D if α · 2r D,
where α is a dimensionless parameter describing the roughness
of a diffuser. Using this approximation, we obtain a simple
estimation of the difference of arrival times t
RD
within A
r
,
t
RD
=
α f
D
C
0
· f
O
· r
if
α f
D
f
O
· r
< 0.5 D
0.5 D
C
0
if
α f
D
f
O
· r
0.5 D
=
1
C
0
· Min
α f
D
f
O
r
, 0.5 D
. (3)
For the out-of-focus point q
shown in Fig. 1 (inset), A
r
corre-
sponds to the area covered by the cone angle θ, and so we have
r
z · θ
NA
n
z and
t
RD
(z) =
1
C
0
· Min
α f
D
f
O
·
NA
n
· z , 0.5 D
. (4)
Combining Eqs. (1) and (4), we finally obtain the effective pulse
duration at an out-of-focus point q
at distance z from the IP,
namely
τ
eff
(z) = τ
0
+ t
RD
+ t
G
(5)
= τ
0
+
Min
α f
D
f
O
NA
n
z , 0.5 D
C
0
+
( f
D
+ f
O
d)NA
2
2 C
0
nf
2
O
z
2
+ n
n
n
2
NA
2
C
0
n
2
NA
2
z, (6)
where τ
0
is the pulse width of the laser source.
Figure 3 shows the numerical results of τ
eff
(z) for the cases of
three different sample objectives commonly used for biomedical
microscopy. Consistent with a previous report,
9
we find that the
contribution of t
G
to τ
eff
(z) is negligible when z is within a few
Rayleigh lengths (z
R
). Nevertheless, in this small z regime, t
RD
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Yu et al.: Wide-field optical sectioning by plane-projection multiphoton microscopy
in Eq. (6) can lead to a significant elongation of pulse width.
In particular, for the small z regions where α f
D
/f
O
· NA/n · z
< 0.5 D,Eq.(6) can be simplified as
τ
eff
τ
0
1 +
α f
D
f
O
·
NA
τ
0
nC
0
z
= τ
0
1 +
α f
D
f
O
·
n λ
πτ
0
C
0
NA
z
,
with
z
z
z
R
π NA
2
n
2
λ
z. (7)
Here,
z is defined in units of Rayleigh length in order to facilitate
the comparison of our results with conventional confocal and
two-photon scanning microscopy. We further define
z
f
O
f
D
·
πτ
0
C
0
NA
n λ
=
πτ
0
C
0
λα f
D
·
f
O
NA
n
, (8)
whereby at
z = z
, τ
eff
2τ
0
, i.e., z = z
R
z
indicates positions
at which the effective pulse width is doubled. For two-photon
excitation, this corresponds to the positions where the fluores-
cence signal drops to half of the maximum. In conventional
confocal and two-photon scanning microscopy, the correspond-
ing
z
1. From the calculations outlined in Fig. 3,wefind
that optical sectioning is comparable with conventional confo-
cal microscopy, with either moderate (0.3–0.75) or high (>1)
NA objectives. Moreover, we find that laser pulses of 100-fs
pulse width are sufficient to provide such sectioning effects.
Fig. 3 Effective pulse width and two-photon excitation strength as a
function of z under different sample objectives. The numerical results
were obtained from Eq. (6). Notice that Eq. (8) predicts
z
3.53,
2.21, and 1.62 for these objectives, respectively, which are comparable
with the numerical results. To estimate the normalized strength of two-
photon excitation S
2p
, the pulse energy was set as a constant and the
inverse of τ
eff
was used to represent S
2p
[Eq. (12)]. The horizontal
(distance) and vertical (τ
eff
) axes are expressed in units of Rayleigh
length and τ
0
, respectively. Parameters: f
D
= 160 mm, D = 100 μm,
α = 1, d = 200 mm, λ = 800 nm, and τ
0
= 100 fs. Sample objective
10×:NA= 0.3, f
O
= 16 mm, n = 1. Sample objective 40×:NA=
0.75, f
O
= 4 mm, n = 1. Sample objective 60×:NA= 1.1, f
O
= 2.67
mm, n = 1.33 (water immersion).
2.2 Efficiency of Multiphoton Excitation at Low
Repetition Rate
To solve the limitation of low frame rate, we next examine
how the repetition rates of pulsed lasers influence the efficiency
of two-photon excitation (at constant average power). In short,
we find that a 10
5
-fold increase in signal-to-noise ratio (SNR) is
obtained by lowering the repetition rate from 100 MHz to 1 kHz,
thus providing a signal level comparable to that of conventional
multiphoton and epifluorescence microscopy.
For simplicity, we consider a two-photon excitation pro-
cess and estimate the light intensity required for wide-field
two-photon excitation. The fluorescence signal obtained from
a single laser pulse at a single pixel is
s
2 p
= β · I
2
p
· τ, (9)
where β is the two-photon excitation coefficient, I
p
is the peak
intensity of the excitation pulse, and τ is pulse width. Within a
time unit, the fluorescence signal from each pixel collected at the
detector (i.e., the CCD camera), S
2p
, depends on the repetition
rate of the pulsed laser f as
S
2 p
= s
2 p
· f. (10)
On the other hand, within a time unit, the average intensity of
the pulsed laser on a single pixel is
I
avg
= τ · I
p
· f. (11)
Here we have assumed that the intensity profile of the pulse is
rectangular in the time domain. Combining Eqs. (9)(11),we
have
S
2 p
= β ·
I
2
avg
f · τ
1
f
, (12)
which suggests that for a fixed average intensity I
avg
, the signal
level can be significantly enhanced by reducing the repetition
rate f. For example, lowering f from 100 MHz to 1 kHz can
increase the signal 10
5
-fold without increasing the average light
intensity delivered to the sample. It should be noted that the I
p
of
our low-repetition-rate setup is of similar order of magnitude as
that used in high-repetition-rate point-scanning microscopies.
Thus, the signal levels of these two schemes are predicted to be
comparable.
3 Methods and Materials
The light sources we used in this work are ultrafast chirped
pulse Ti:sapphire amplifiers. Two different models were used
for the travel convenience of the authors. Live-cell imaging was
studied (see Figs. 5 and 6 and Video 1) with a Spectra-Physics
R
Spitfire
R
Pro, seeded with a Spectra-Physics
R
Mai Tai
R
SP
ultrafast oscillator situated parallel to the amplifier within an
enclosure. Quantitative characterizations of our technique (see
Figs. 4, 7, 8,and9) were carried out with a Coherent
R
Legend
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Yu et al.: Wide-field optical sectioning by plane-projection multiphoton microscopy
Fig. 4 Theoretical (gray line) and measured (black dots) fluorescence
signal from a homogeneous dye film. The axial resolution was defined
as the FWHM of the fluorescence profile. The fluorescence profile was
obtained by taking optical sections of a homogeneous dye film (thick-
ness less than 2 μm). The signal was determined by the integrating the
intensity of each section. The profile was normalized by its maximum.
The FWHM here is 2 μm, while the theoretical axial resolution of
a confocal microscope with the same objective is 0.6 μm. Parame-
ters in theoretical estimation and experiment: M
L
O
= 60, NA = 1.42,
n = 1.5 (oil immersion), f
O
= 160/60 mm, f
D
= 160 mm, τ
0
= 35 fs,
d = 200 mm, λ = 800 nm, α = 0.1, D = 30μm.
Elite-USP-1k-HE, seeded with a Coherent
R
Mantis-5 ultrafast
oscillator located parallel to the amplifier. The full width at
half maximum (FWHM) pulse duration of both amplifiers was
approximately 35 fs or less. The wavelength of both amplifiers
was centered approximately at 800 nm with FWHM 30 nm. We
expanded the beam size by a lens pair, such that the beam profile
on the diffuser was two-dimensional (2D) Gaussian with FWHM
10 mm. The maximal output of the laser amplifier was 3W
(average power), and was attenuated to avoid thermal damage to
biological samples. The average laser powers reported in Secs. 4
and 5 were all measured at the back aperture of the microscope
objective L
O
.
The optical diffuser employed was a Thor Labs model DG10-
120. Diffusers, in general, can cause significant inhomogeneities
of the light intensity at the image plane. To reduce these inhomo-
geneities, glass etching cream (Armour Etch
R
) was used to etch
the diffuser. The roughness parameters D and α of the diffuser
were found to be 30 μm and 0.1 after etching, according to the
surface profile we measured (data not shown).
AsshowninFig.1, the collimated laser beam is scat-
tered by the optical diffuser, collected by the diffuser lens
L
D
, and then projected to the sample via the sample objectives
(LUMFLN 60×WNA1.1,PLANAPON60×O NA 1.42, and
UPLFLN 10×NA 0.3, Olympus). The LUMFLN model objec-
tive was used for the living biological samples owing to its long
working distance. The PLANAPO objective was used for the
quantitative characterizations and the fixed biological sample.
The UPLFLN objective was used for demonstration of large
fields of view. Because our theoretical model is based on ge-
ometrical optics, it is essential that L
D
has sufficient NA to
geometrically resolve the scattering structures. In practice, L
D
with f
D
= 160 mm and NA 0.2 was chosen.
The chromatic dispersion of the full optical path was precom-
pensated by the built-in compressor of the ultrafast amplifier(s)
such that the signal level at the IP was optimized. The image was
obtained by a CCD camera (iXon DU-885K, Andor) through L
I
.
The field of view is square and of length 6.4 mm/M
L
O
per
side, where M
L
O
is the nominal magnification of L
O
. The illu-
mination fi eld is 2D Gaussian with FWHM 10 mm/M
L
O
.A
larger illumination field or more uniform profile can be obtained
by further expanding the laser beam before the optical diffuser.
The axial resolution was determined by taking images along
the optical axis of a homogeneous film (thickness less than 2
μm) of fluorescent dye (F-1300, Invitrogen). For live cell imag-
ing, we used human mammary gland MCF-10A cells express-
ing cyan fluorescent protein-conjugated histone (H2B-cerulean),
which binds to chromosomes and has been widely used to indi-
cate cell nuclei. MCF-10A cells were seeded in 3D matrigel (BD
Matrigel
TM
) for 10 days to form bowl-shape cell clusters several
hundred micrometers in size. We then used the cell clusters to
evaluate the high-frame-rate acquisition and optical sectioning
capabilities of our PPMP microscope. Following the acquisi-
tion of optical sections, 3D views of the epithelial tissue were
reconstructed using the ImageJ 3D Viewer.
To compare the imaging speed of our setup with that
achieved by a temporal focusing microscope using a nonam-
plified oscillator,
9
we imaged dye-stained Mardin Darby canine
kidney (MDCK) cells. The cells were cultured on glass for
2 days, followed by fixation with 4% paraformaldehyde and
staining with Hoechst 33342 (a chromosome-staining dye from
Invitrogen, comparable in brightness to the DAPI dye used in
Ref. 9).
4Results
4.1 Axial Resolution of Plane-Projection
Multiphoton Microscopy is Comparable to
Conventional Confocal Microscopy
Figure 4 shows the axial resolution of the optical setup depicted
in Fig. 1. T he axial resolution was determined by the FWHM of
the fluorescence signal. With M
L
O
= 60, NA = 1.4, n = 1.5, the
axial resolution was found to be 2 μm, and the corresponding
z
3. This is comparable to the axial resolution of an op-
timized conventional confocal microscope, which has
z
1.
Note that it should be possible to obtain an axial resolution of
z
1 by optimizing the microscope design, as we described in
Sec. 5.
4.2 Frame Rate of Plane-Projection Multiphoton
Microscopy is Comparable to Conventional
Epifluorescence Microscopy for Live-Tissue
Imaging
To demonstrate that PPMP microscopy has the capability of
imaging live tissues at high frame rate, we performed optical
sectioning of live, three-dimensional MCF-10A cell clusters
having hemispherical shapes (Fig. 5). The images were then
used to reconstruct the 3D view (Fig. 6 and Video 1). Here, the
exposure time was set at 200 ms, equivalent to 5 fps, which is
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Yu et al.: Wide-field optical sectioning by plane-projection multiphoton microscopy
Fig. 5 Optical sections and lateral view of live MCF-10A cells in a
hemispherical structure. The top panel shows the sections at different
depths. The middle and bottom panels show the reconstructed lateral
views under a PPMP microscope and an epifluorescence microscope
(Epi), respectively. In the lateral view from the epifluorescence micro-
scope, we clearly observe the residual out-of-focus light at the top
and bottom edges of the nuclei. The lines in the middle of the top
views indicate the positions where the lateral views were taken. Flu-
orescence signals were from cell nuclei expressing cyan fluorescent
protein-conjugated histone (H2B-cerulean), which binds to chromo-
somes. Exposure time of each frame: 0.2 s. L
O
:60×,NA= 1.1, n
= 1.33. Step size: 1 μm. Laser average power: <10 mW.
10 times faster than the conventional multiphoton microscope
we also used to image the same sample, and is 150 times faster
than a temporal focusing setup using a nonamplified 75-MHz
Ti:sapphire oscillator to image cells stained with (much brighter)
fluorescent dye.
9
Such an exposure time also lies within the same
order of magnitude of that typically used in conventional epiflu-
orescence microscopy (100 ms), which we also used to image
the same sample (Fig. 5, lower panel) through the same CCD
camera with a mercury vapor lamp (X-Cite
R
120Q, Lumen
Dynamics, attenuated by OD 2 to prevent significant photo-
bleaching). Figure 5 demonstrates resolution of the boundaries
of cell nuclei along the z axis (i.e., the optical axis). Nuclei in
the PPMP image appear oval in structure, resembling the normal
shape of cell nuclei. In contrast, the lateral view obtained from
epifluorescence microscopy (Fig. 5, lower panel) shows distor-
tion of the proper cell nuclear shape, due to the spreading of the
out-of-focus signal in an epifluorescence microscope. These re-
Video 1 Reconstructed 3D view of the live MCF-10A cells in Fig. 5.
The 3D view was reconstructed from 40 sections with a 1-μmstep
size, using ImageJ 3D Viewer (QuickTime, 2.1 MB).
[URL: http://dx.doi.org/10.1117/1.3647570.1]
sults suggest that PPMP microscopy can obtain high-frame-rate
optical sectioning on live tissues.
4.3 Inhomogeneity of the Illumination Field Can be
Reduced by Rotating the Diffuser
In this study, we found that conventional diffusers can cause a
significant inhomogeneity of the light intensity in the illumina-
tion field, i.e., bright spots [Fig. 7(a)]. Although optical diffusers
are known for generating speckle when used with temporally and
spatially coherent light sources,
10
we believe that the observed
field inhomogeneity is not from speckles, but rather is an effect
of the diffuser used.
One way to examine the physical basis of the field inhomo-
geneity is to compare the images of a (homogeneous) fluorescent
film with the diffuser at the focus and at a position slightly out
of focus (i.e., with the diffuser translated along the optical axis
by a short distance). If the bright spots are speckles, slightly
defocusing the diffuser will only rearrange the location of the
speckles, while the statistical properties such as the mean and
the variation of brightness will remain the same.
10
The results
of this experiment are shown in Fig. 7. After the diffuser was
translated 3 mm from its in-focus position, the bright spots ap-
pearing in the in-focus image [Fig. 7(a)] were found to remain in
place and become dimmer and blurred [Fig. 7(b)]. These mea-
surements suggest that the observed bright spots are not from the
speckle effect, but rather from the randomness of the scattering
structures distributed on the diffuser surface.
The observed field inhomogeneity also leads to inhomoge-
neous sectioning across the field of view; the level of which can
be measured by imaging a homogeneous dye film, then sepa-
rating the field of view into several areas and comparing the
FWHMs of their intensity profiles. In our setup, the standard
deviation of the FWHMs was found to be 0.3 μm. One way
to reduce this inhomogeneity is through the use of multiple dif-
fusers. However, each diffuser would generate a certain level of
time delay and thus contribute to pulse broadening. At the image
plane, the telescoping can only recover the pulse width at the
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Yu et al.: Wide-field optical sectioning by plane-projection multiphoton microscopy
Fig. 6 Reconstructed 3D view of the live MCF-10A cells in Fig. 5.The
3D view was reconstructed from 40 sections with a 1-μm step size
using ImageJ 3-D Viewer. Note that the boundaries of cell nuclei along
the z axis (i.e., the optical axis) can be clearly observed; a capability
that cannot be achieved by conventional epifluorescence microscopy.
input to the last diffuser (after it has already been broadened by
the first diffuser). Thus, having multiple tandem diffusers can
reduce the axial resolution of the optical sectioning. One way
to mitigate this effect is by proper ordering of the diffusers. As
mentioned in Sec. 2, diffusers with larger D and α induce longer
random time delay. Therefore, in such multiple-diffuser setups,
the diffuser with smaller D and α should be placed before the one
with larger D and α. As an alternative solution, we have chosen
to simply rotate the diffuser. By rotating the optical diffuser dur-
ing the acquisition of a single frame, the inhomogeneities in the
illumination field are averaged out. This effect is demonstrated
in Fig. 8.
5 Discussion
5.1 Optimization and Limit of Axial Resolution
Here we discuss the further optimization of PPMP microscopy,
and the axial resolution limit that can ultimately be obtained,
Fig. 7 Images of illumination field inhomogeneity with the optical dif-
fuser (a) in-focus and (b) out-of-focus. For the out-of-focus image, the
optical diffuser was translated for 3 mm along the optical axis from
the in-focus position. Here we can see that the out-of-focus image does
not possess the same magnitude of field inhomogeneity as the in-focus
image. However, while blurred, the locations of the bright spots are
preserved in the out-of-focus image, consistent with the field inhomo-
geneity resulting not from speckle, but rather from the randomness of
the scattering structures on the surface of the diffuser. The sample is
a homogeneous dye film. The dimension of the field shown here is
70 × 70 pixels.
as derived within the realm of geometrical optics. Equation (8)
suggests that
z
can be further reduced by using an objective with
a higher magnification and NA (which often exhibits a smaller
f
O
NA/n),asshowninFig.3. Likewise, increasing f
D
, α,or
reducing τ
0
leads to smaller z
. We stress that these estimations
are derived based on geometrical optics, and so may not be valid
in the extreme case where
z
< 1, in which case the optimal axial
resolution of our temporal focusing setup would be the same as
that of a laser scanning microscope.
9
A further fundamental advantage of diffuser-based temporal
focusing over blazed grating approaches is that the diffuser-
based technique can achieve the axial resolution of a point-
scan setup, whereas (single) grating-based temporal focusing is
limited to that of a line-scan setup. The difference arises from the
way in which the time delays are generated. For optical diffusers,
the time delay results from the surface roughness of the diffusers,
which creates a 2D spatial profile for the randomness of the
time delay. In contrast, the time delay in grating-based temporal
focusing is created by the one-dimensional scan of the laser
pulses on the grating surface; such a mechanism restricts the time
Fig. 8 Illumination field intensity inhomogeneity with (a) fixed and (b)
rotated optical diffusers. The field inhomogeneity is defined as the stan-
dard deviation of the field divided by the average intensity of the field.
The field inhomogeneity is greatly reduced by rotating the optical dif-
fuser during the exposure of each frame. The sample is a homogeneous
dye film. The dimension of the field shown here is 100×100 pixels.
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Yu et al.: Wide-field optical sectioning by plane-projection multiphoton microscopy
delay to be one-dimensional. This restriction has been overcome
by using two orthogonally aligned gratings.
11
In such a setup,
the two gratings must differ in groove density sufficiently, such
that the scanning of the laser pulse can be well separated in two
orthogonal dimensions. Such a design increases the complexity
of the apparatus and will likely require multiple pairs o f gratings
when multiple/tunable excitation wavelengths are used.
From Eq. (4), the spread, or distribution, of arrival times pro-
duced from the surface roughness of a diffuser is upper bounded
by the factor D. This suggests that diffusers with larger D should
be used to ensure a sufficiently large spread of arrival times.
The roughness of the diffuser surface, however, leads in turn
to roughness of the image plane, D
. Using the thin lens for-
mula, we estimate D
to be (f
O
/f
D
)
2
D. This suggests that D
can
be negligible if f
D
f
O
. Thus, with a proper arrangement of
parameters, the roughness of the image plane can be reduced
below one Rayleigh length, while the surface roughness of the
diffuser is sufficiently large to create temporal focusing.
5.2 Limitation of Frame Rate and Benefits of Low
Repetition Rate
For live-tissue imaging, the frame rates of our setup are limited
by the low SNR of fluorescent proteins expressed in living sys-
tems. Nevertheless, Eq. (12) suggests that SNR can be further
enhanced by lowering the repetition rate while maintaining the
average power of the laser. For example, the frame rate of our
setup can be further increased by equipping our system with a
pulsed laser of much lower repetition rate, e.g., 100 Hz. With
such a low repetition rate, Eq. (12) suggests a 10-fold stronger
SNR than what is presented in this study. This would lead to a
frame rate of up to 50 fps, a rate sufficient to study most biolog-
ical processes such as cell division, migration, and polarization.
This limiting frame rate is estimated for imaging the flu-
orescent proteins expressed in living systems. This limitation
is relaxed, though, if the signals are derived from materials
with strong fluorescence efficiency such as fluorescent dyes and
nanoparticles. In such cases, we are able to achieve higher frame
rates, as shown in the top panel of Fig. 9, which was obtained
with an imaging speed 1000 times faster than a temporal focus-
ing setup using a nonamplified 75 MHz Ti:sapphire oscillator
and a similarly dyed sample.
9
Furthermore, by using microscope
objectives of low magnification, we can also achieve large fields
of view. This capability for large field of view is demonstrated
in the bottom panel of Fig. 9, where we imaged dye-stained
MDCK cells with a 10× objective. The size of the field of
view shown here is 0.6 × 0.6 mm
2
and the exposure time was
100 ms.
Our setup can achieve these larger fields of view with a
relatively short exposure duration simply because the 1-kHz
amplifier is very powerful; that is, because it is supplying its
average power at a low repetition rate and low duty cycle and thus
achieving a high peak power. To generate multiphoton excitation
at the level required for imaging with reasonable frame rates, the
peak intensity is commonly around or greater than 1 kW/μm
2
.
12
Therefore, to excite an area up to 1 mm
2
, one needs a light source
with peak power greater than 10
9
W. The maximal peak power of
our amplifier is roughly 10
11
W, and is thus powerful enough to
support a large field of view for most microscopy applications.
It should be noted that in the original temporal focusing setup,
9
Fig. 9 MDCK cells stained with fluorescent dye (left) and under bright-
field illumination (right) using 60× (top) and 10× (bottom) microscope
objectives. The imaging speed using the 60× objective is 1000 times
faster than a temporal focusing setup using a non-amplified 75 MHz
Ti:sapphire oscillator and similarly dye-stained sample (Ref. 9). Signal
from areas outside of the nuclei results from non–specific staining.
Exposure time: 0.03 (60×) and 0.1 (10×) seconds. L
O
:60×/10×,NA
= 1.42/0.3, n = 1.5/1. Laser average power: <15/60 mW.
a 140 × 140–μm
2
field of view was obtained with an average
power of 30 mW and an exposure time of 30 s. This indicates
that a 1-mm
2
field of view can be achieved with that instrument
by using a low magnification objective and an average power of
around 1.5 W, though the exposure time in such a setup could be
slightly longer than 30 s because lower magnification objectives
are typically less efficient in collecting light.
However, from a biologist’s point of view, we would also
like to point out that discussing the imaging speed for fixed
biological samples stained with fluorescent dye is less important
than the speed achievable for living systems. Once a sample
is fixed, using an imaging time of either 3 h or 10 s would
most likely provide the same level of details and information.
On the o ther hand, for studies of dynamic biological processes,
the imaging speed would determine the temporal resolution of
the observations. To the best of our knowledge, this is the first
report of imaging live cells expressing fluorescent protein by a
temporal focusing microscope at a frame rate faster than 1 fps.
In addition to the enhancement of the signal level and frame
rate, there are certain potential benefits provided by lowering
the repetition rate from the megahertz to kilohertz regime. It
has been reported that the use of low repetition rates (at the
same optical power) can reduce photobleaching.
13, 14
This is
achieved through the avoidance of dark state conversion. Indeed,
a 5- to 25-fold enhancement of total fluorescence yield, before
detrimental effects from photobleaching, has been reported.
13
Moreover, lowering the repetition rate is equivalent to providing
the system a longer window of no excitation. This would allow
slow processes such as heat dissipation to occur more efficiently,
thus minimizing sample damage caused by a continuous accu-
mulation of heat.
4
As a result, even with a similar amount of
thermal energy introduced by the excitation process, a sample
excited at a lower repetition rate is less likely to be damaged by
heat accumulation as compared to the use of a higher repetition
rate.
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Yu et al.: Wide-field optical sectioning by plane-projection multiphoton microscopy
5.3 Potential Applications as Structured Light
Microscopy
In principle, the inhomogeneity of the illumination field can
be utilized for structured light microscopy.
6
This could be par-
ticularly useful in applications where reasonable optical sec-
tioning, as provided by temporal focusing, is not achievable.
Examples include coherent anti-Stokes Raman scattering
(CARS) and stimulated Raman scattering microscopy, where
picosecond pulses are generally required to obtain chemical
specificity.
15, 16
BasedonEq.(8), pulse widths of picosecond
duration would greatly reduce the sectioning effect. Neverthe-
less, by using the inhomogeneity of the illumination field as a
structured light source, it is possible to regain the optical sec-
tioning of these systems, as demonstrated in a previous study.
6
This allows one to integrate CARS with multiphoton excitation
in a wide-field microscope simply by using an optical diffuser.
6 Conclusion
The question of how to increase image acquisition rate and axial
resolution, while maintaining a biocompatible laser dosage, is a
long-standing challenge in the optical microscopy community.
In this report, we have demonstrated a microscope design for
live-tissue imaging that provides an axial resolution compara-
ble to confocal microscopy and a frame rate similar to that of
epifluorescence microscopy.
By utilizing an optical diffuser, a temporal focusing setup
is realized with a design as simple as a conventional epiflu-
orescence microscope. Even at a high frame rate, the photo-
bleaching and thermal damage of PPMP microscopy could be
lower than conventional multiphoton and confocal laser scan-
ning microscopy. Compared with temporal focusing techniques
using Megahertz repetition-rate laser pulse trains, the use of
low repetition-rate pulses, while maintaining the same average
power, can significantly enhance the SNR. In addition, using
an optical diffuser instead of a blazed diffraction grating pro-
vides flexibility for multi- or tunable-wavelength light sources,
and thus creates a platform for multispectral imaging and
pump-probe microscopy. Taken together, these features sug-
gest that plane-projection multiphoton microscopy can be used
to study fast, three-dimensional processes in living cells and
tissues, and to do so with minimal phototoxicity and thermal
damage.
Acknowledgments
This work is a collaborative research effort of the Cell Polar-
ity Laboratory at the California Institute of Technology and the
Technology and Applications Center at Newport Corporation.
The authors would like to thank Mr. Craig Goldberg of New-
port Corporation for facilitating the collaboration, and Professor
Shi-Wei Chu for the inspiring discussions. Mr. Ji Hun Kim of
Caltech is sincerely acknowledged for the measurement of the
surface profile of the optical diffuser. CLG recognizes support
from the Ellison Medical Foundation and the Weston Havens
Foundation. Support from the National Science Foundation
Chemistry Research Instrumentation and Facilities: Instrument
Development program to GAB is gratefully acknowledged.
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