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Departments of Physiology and Medicine, University of Maryland School of Medicine, Baltimore, Maryland 21201
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ABSTRACT |
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 TOP
 ABSTRACT
 INTRODUCTION
METHODS AND MATERIALS
RESULTS
DISCUSSION
REFERENCES
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We describe a custom one-photon (confocal) and two-photon all-digital (photon counting) laser scanning microscope. The confocal component uses two avalanche photodiodes (APDs) as the fluorescence detector to achieve high sensitivity and to overcome the limited photon counting rate of a single APD (~5 MHz). The confocal component is approximately nine times more efficient than our commercial confocal microscope (fluorophore fluo 4). Switching from one-photon to two-photon excitation mode (Ti:sapphire laser) is accomplished by moving a single mirror beneath the objective lens. The pulse from the Ti:sapphire laser is 109 fs in duration at the specimen plane, and average power is ~5 mW. Two-photon excited fluorescence is detected by a fast photomultiplier tube. With a ×63 1.4 NA oil-immersion objective, the resolution of the confocal system is 0.25 µm laterally and 0.52 µm axially. For the two-photon system, the corresponding values are 0.28 and 0.82 µm. The system is advantageous when excitation intensity must be limited, when fluorescence is low, or when thick, scattering specimens are being studied (with two-photon excitation).
cardiac muscle; fluorescence; excitation-contraction coupling; microscopy
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INTRODUCTION |
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TOP
ABSTRACT
INTRODUCTION
METHODS AND MATERIALS
RESULTS
DISCUSSION
REFERENCES
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OPTICAL SECTIONING TECHNIQUES, such as confocal
microscopy (14) and two-photon microscopy (6), provide the means to
localize molecules in living cells with high spatial and temporal
resolution. With these techniques, the understanding of
cell function is being extended to cellular volumes
("microdomains") on the order of 10
15 liters (16). Furthermore, events
within such microdomains of individual cells may be visualized as part
of a much larger scene, such as an entire group of interconnected
cells. Through the use of two-photon microscopy, such
scenes may be located deep (e.g., ~100 µm) within a living,
reasonably intact tissue (3). The physiologist studying live tissue in
this situation is usually faced, however, with measuring low levels of
light. It may be necessary to use low illumination intensity to reduce
photodamage to the specimen (12). There may be only a few fluorescent
molecules, or the concentration of fluorescent dye must be kept low to
reduce perturbation of the cell (2). In a confocal system, spatial resolution can be improved (up to a point) by rejecting fluorescence, further reducing the signal. In thick specimens, light scattering may
reduce the efficiency of fluorescence excitation and detection (3).
Because we are interested in conducting studies of the type described
above, we have constructed a combined confocal and two-photon
microscope in which the unique features are related to efficient
detection of low levels of fluorescence. This is achieved through
simple, efficient optical design, the use of novel detectors (for such
microscopes), and photon counting, which offers improved accuracy and
signal-to-noise ratio (S/N) at low light levels (compared with analog
methods). The use of photon counting also allows the system to be
"all digital," with the consequent advantages of accuracy and
ease of signal processing. Thus the microscope is a "photometric"
device, in that light is quantified accurately, for the purposes of
quantification of fluorescent molecules. We also present a simple
technique for measuring duration of the ultrashort pulse of infrared
(IR) light used for two-photon excitation at the specimen plane. This
technique allowed us to measure the group velocity dispersion (8, 15)
introduced by the glass in the microscope.
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METHODS AND MATERIALS |
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TOP
ABSTRACT
INTRODUCTION
METHODS AND MATERIALS
RESULTS
DISCUSSION
REFERENCES
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Optical components.
Figure 1 presents a schematic diagram of
the optical components and the fluorescence detectors. The major
optical component is a Nikon Diaphot 300 microscope (Nikon, Garden
City, NY) with two Nikon objective lenses (L4), an oil-immersion
×63 lens (NA 1.4) and a ×60 "water objective" lens
(NA 1.2). The microscope and associated optics and lasers are all
mounted on a large vibration isolation table (4 × 6 ft)
(Technical Manufacturing, Peabody, MA) with the "clean top"
surface provided with mounting holes. Most optical components (holders
and mirrors) were purchased from New Focus (Sunnyvale, CA). The design
of the confocal laser "scan side" of the instrument is
conventional and has been described previously (13). Differences to the
system, also described previously, include the method of beam scanning,
beam intensity modulation, fluorescence detection, and digital image
construction. On the "confocal side," the beam from an argon ion
laser is incident on a single mirror (SM), located in the exit pupil of
the scan lens (L3). This mirror is rotated by galvanometers around two orthogonal axes (x, horizontal or line scanning; y,
vertical or frame scanning). The line-scanning galvanometer (model
6800, Cambridge Technology, Watertown, MA) is mounted in a custom,
inertially balanced holder mounted on the shaft of the large (frame
scanning) galvanometer (model 6900, Cambridge Technology). This system
offers several advantages over dual-mirror systems in that the design is simplified, light losses are minimized, and alignment is simplified. The intensity of the beam is modulated using an electrooptic device (i.e., a Pockels cell, model 350-50, Conoptics Danbury, CT). This device provides "blanking" of the beam during horizontal and
vertical "flyback" of the scan mirror. In addition, it provides
precise, programmed control of beam intensity, which is distinctly
advantageous compared with the use of conventional neutral-density
optical filters mounted in wheels. The L3 scan lens (Fig. 1) is a Nikon ocular (×10, wide-field CFW) adapted to fit into the side video port of the microscope. Motorized focusing
(z-axis) is accomplished with a Nikon Remote focus accessory
(stepper motor) controlled through a serial computer interface (RS
232). Fluorescence emission from the specimen is "descanned" by
the scan mirror, and, after passing back through appropriate dichroic
beam splitters (EFSP1 and DCLP1) and filters (BPF2), is directed by a
mirror (M1) into an aperture. This aperture is relatively large
(~1.0-mm diameter) because the Airy disc (i.e., the image of the
point source of emitted light) is imaged "at infinity." Light
passing through the aperture is essentially parallel and is focused by
a lens (L1) into the detection system. This system consists of a
50%-50% beam splitter (BS1), which directs light equally into
two photon-counting modules, each consisting of an integrated avalanche
photodiode (APD) and amplifier/discriminator (SPCM-AQ-121; EG&G
Optoelectronics). The L1 lens focuses the light to a region smaller
than the diameter of the active area of the photodiode (~180 µm).
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Electronic components. The essential electronic components of the microscope are 1) a microcomputer equipped with a variable-scan digital video capture board, a digital input/output (I/O) board, and a digital-to-analog (D/A) output board; 2) a custom pulse counter that emulates a variable-scan digital video camera; and 3) driver circuitry for the mirror galvanometers. Control of most functions is through a "virtual instrument" consisting of a LabView program (LabView and IMAQ, National Instruments) operating the digital video capture board, D/A board, and I/O board in the microcomputer (operating system Windows NT 4.0). The digital I/O board generates horizontal and vertical synchronization pulses and a pixel clock signal, all of which are used to gate the counting circuitry and to synchronize capture of the digital data by the variable-scan video capture board. Voltage waveforms, representing the desired scan mirror excursions, are generated by the D/A board. Actual driving voltage is derived from these signals by driver circuitry provided with the galvanometers (Cambridge Technology). Pulses (transistor-to-transistor logic) from the detectors (APDs and PMT) are counted by a custom counter (gated by the pixel clock) and sent as 8-bit digital data to the variable-scan digital video capture board. The custom gated counter has a maximum count rate of 50 MHz per channel. It has two separate channels that each produce a number (0-255) representing the counts detected during a pixel. The numbers from the two channels can be sent separately to the digital video capture board, or they can be added to each other before being sent. This provides the capability for dual-channel imaging (confocal) or the use of two APDs to improve S/N (see below). Images are simultaneously displayed and stored in real time (either in the computer RAM or on a hard disk). In addition, a D/A circuit within the custom counter produces analog voltage signals that are viewed on an oscilloscope and that could be used with an analog variable-scan video capture board.
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RESULTS |
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TOP
ABSTRACT
INTRODUCTION
METHODS AND MATERIALS
RESULTS
DISCUSSION
REFERENCES
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System calibration and characteristics: low light detection. APDs are advantageous for measuring low levels of light, primarily because they have higher quantum efficiency (QE) than PMTs at the wavelengths of interest. They are also ideal for confocal microscopy, where the emitted light is descanned and can be focused to a small point, corresponding to the small active area of an APD. Furthermore, APDs with particularly small active areas with consequent low noise can be utilized. The QE of the APDs we used increases from 0.45 to 0.60 over the wavelengths from 500 to 600 nm (range of the emission spectrum of Ca-fluo 4), whereas the QE of an appropriate PMT falls from ~0.15 to ~0.05 over that same range. To estimate the improved detection of fluo 4 fluorescence that might be provided by an APD compared with a PMT, we integrated the numerical product of the fluorescence emission spectrum of Ca-fluo 4 and the QE of our APDs and PMT. The APD produced 4.2 times as much signal from fluo 4 as did the PMT.
Photon counting, which can be done with either APDs or the PMT, is particularly advantageous when levels of light are low (1). Photon counting eliminates the variability arising from the distribution of current pulse amplitudes and provides an absolutely accurate quantification of "black level," which arises only from "dark counts." The latter is particularly advantageous when using a low background fluorescence (F0) to construct ratioed fluorescence signals, such as those used to study cardiac Ca2+ sparks (4, 10). The only significant disadvantage is that maximum count rates in photon-counting systems are limited by system "dead times," or that time during which the system is insensitive to another incoming photon. Thus it is necessary to correct the recorded count rates for system dead time when count rates are high. The correction factor required for a particular APD is provided by the manufacturer. However, the use of a correction factor increases uncertainty (reduces S/N) of the measurement, and therefore correction factors should be kept as low as possible. The photon-counting modules we used had a dead time of ~40 × 109 s, and the corrected count rate as a
function of measured count rate is shown in Fig.
2A (filled symbols). It can be seen
that the corrected count rates become appreciably different (i.e., >10%) for count rates >2.0 MHz. This problem can be partly
overcome through the use of two APDs such that the count rate at each
is 50% of what it would be otherwise (Fig. 2A, open symbols).
This reduces the correction factor that must be applied. Thus the
correction factor for a two-APD detector at a recorded count rate of
10.0 MHz is 1.37, whereas that for a single APD detector is 2.0. The use of correction factors is perfectly valid, provided that a good
estimate of the measured count rate is obtained. This requires that
pixel durations be long enough that adequate numbers of counts are
accumulated. For example, in a single-APD system, count rates of 5.0 MHz can be measured accurately with pixel durations of 10.0 µs, but
not 1.0 µs. With 10.0-µs pixels, each pixel will contain, on
average, 50 counts. The variance of a sample of such pixels
(representing, for example, the peaks of Ca2+ sparks) will
be 7.07, corresponding to 0.707 MHz. With pixel durations of 1.0 µs,
however, the variance will be 2.24, corresponding to 2.24 MHz. Thus
much larger and more variable correction factors will be used for
pixels 1.0 µs in duration compared with pixels of 10.0 µs in
duration, despite the fact that the "true" count rate (i.e., 5.0 MHz) is exactly the same. As shown in Fig. 2A, this problem is
reduced significantly by the use of two APDs.
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Two-photon excitation.
Efficient two-photon excitation requires ultrashort pulses of light at
the specimen plane (5). It is known, however, that the duration of the
pulse of light increases as it travels through glass, such as
microscope objectives (8). The amount of such broadening can be
calculated, provided that the amount of glass and its refractive index
are known. This is generally not known for commercial microscopes and
objective lenses, however, so a method of determining "pulse
width" at the specimen plane was devised. The ability to determine
pulse width accurately at the specimen plane also aids in compensating
GVD through the use of external "precompensation optics." In our
system, precompensation for GVD is provided by a pair of high-index
prisms (SF- 10) mounted in micromanipulators, according to methods
described earlier (7). The polarization of the output beam is rotated
90° through the use of a periscope (MP1) before precompensation.
Pulse width was determined from the autocorrelation function (Fig.
2D) measured at the specimen plane by using an aspheric lens
(New Focus ×60, NA 0.6) to collimate the light emerging from the
objective lens (Nikon ×60 oil, NA 1.4) before directing it into
an autocorrelator (SP 409-08, Spectra Physics). The aspheric lens added
a negligible amount of glass to the system. We found that this
arrangement was easier to use than the methods described previously (8, 15), in which the GVD compensation is inferred from the maximization of
two-photon fluorescence emission. With 86.5 cm of air between the
prisms, and with 1.8 cm of prism glass in the path, the pulse width at
the specimen plane was 109 × 1015
s. The total compensation was thus 4,075 fs2, similar to
that reported by others (8, 15). The spectral output was approximately
Gaussian, centered at 800 nm, with an FWHM of 11 nm, as determined with
the monochromator and video camera system (Fig. 1).
Spatial resolution.
Point-spread functions (PSF; Fig. 3) were
measured with the use of subresolution fluorescent beads (0.1 µm in
diameter) immobilized on a glass coverslip that forms the bottom of our
standard physiological recording chamber. A three-dimensional data set
was acquired by first focusing on a plane immediately above the bead
and then scanning with the automatic stepper set in 0.1-µm steps.
With a ×60 1.4-NA oil-immersion objective, the lateral FWHM of
the confocal PSF was 0.25 µm and the axial FWHM was 0.52 µm. For
the two-photon system, with the same objective lens, the corresponding values were 0.28 and 0.82 µm. For these measurements, pixel
dimensions were 0.05 µm laterally and 0.1 µm vertically.
Three-dimensional images of the PSF were constructed, and vertical
sections (X-Z planes) are shown in Fig. 3.
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Efficiency.
We compared the optical efficiency of our custom confocal with that of
a Bio-Rad MRC 600 confocal microscope (Bio-Rad Laboratories, Melville,
NY). Efficiency will be defined here as the ratio of the number of
photons exciting a sample (per pixel) to the number detected (per
pixel). The sample was a uniform solution of fluorescein (1.0 µM) in
a physiological recording chamber, and the chamber, containing exactly
the same sample, was used with both microscopes. The MRC 600 was put
into photon-counting mode and adjusted as described in the supplied
manual, with the exception that the gain was set to the maximum.
Excitation power, at the specimen plane, was measured with a power
meter (LaserCheck; Coherent, Auburn, CA) having a resolution of 0.01 µW. The MRC 600 was set to normal scan mode [768 × 512 pixels, 0.275 µm/pixel, 1.5 µs/pixel (estimated)]. Power at
the specimen plane was 5.5 µW (or µJ/s), corresponding to 2.16 × 107 photons/pixel (there are 2.46 × 1018 photons per joule at a wavelength of 488 nm). The mean
number of detected photons per pixel was 2.8, giving an actual
efficiency of 1.30 × 107. For the
custom confocal microscope, a power of 0.97 µW was used to produce
26.7 photons/pixel (mean) in a smaller image (256 × 256 pixels,
0.10 µm/pixel, 10.0 µs/pixel). The efficiency of this system was
thus 1.12 × 106. Thus the
efficiency of the custom confocal is ~ 8.6 times that of the MRC 600. Because the spatial resolution of our system is somewhat better than
that of the MRC-600 (13), it seems reasonable that the relative
efficiency would be even greater if the systems were compared at
exactly the same spatial resolution.
Confocal images of cardiac sparks.
We are interested in recording Ca2+ sparks in cardiac
cells. For this application, the confocal microscope is the clear
choice, because it provides superior spatial resolution to the
two-photon microscope. However, no measurements of Ca2+
sparks with the two-APD system have been published previously. The
preparation of rat cardiac ventricular cells for this purpose was
according to the methods described previously (10). Ca2+
sparks were recorded in the "line-scan" mode, in which we scanned a single line 25.6 µm in length in 3.0 ms. Of this time, 0.44 ms are
used for horizontal flyback. Thus the pixel "dimensions" are 10.0 µs and 0.1 µm. As discussed previously, the relatively long pixel
duration limits the distance that can be scanned but increases S/N. The
small pixel size of 0.1 µm is necessary to satisfy the Nyquist
criterion, given the dimensions of the confocal PSF. Through the use of
the Pockels cell, the intensity of emitted fluorescence in the cardiac
cell at rest (F0) can be adjusted within the range of
1.0-2.0 MHz. Because the peak fluorescence pseudoratios
(
F/F0) of cardiac Ca2+ sparks seldom exceed
4 or 5, this permits the measurement of Ca2+ sparks with a
minimal correction factor and adequate S/N. An example of a
Ca2+ spark with a peak F/F0 of ~2 is shown
in Fig. 4A. The distribution of
amplitudes of all Ca2+ sparks recorded in this experiment
is shown in Fig. 4C (open bars). According to theory,
Ca2+ spark amplitude histograms are limited at their lower
end by noise (9). This limitation was characterized by the distribution of peak F/F0 (Fig. 4C, hatched bars) obtained
from images in which no Ca2+ spark occurred. Clearly,
Ca2+ sparks having F/F0 <0.5 cannot be
measured with this system because of noise. Ca2+ sparks
recorded with two-photon excitation (Fig. 4B) can appear similar to those recorded with one-photon excitation. The amplitude histogram (Fig. 4D), however, is skewed to larger
Ca2+ sparks because greater noise (at the level of
illumination used) limited the ability to measure small
Ca2+ sparks.
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Two-photon images.
Two-photon imaging is expected to be advantageous, particularly in
thick, scattering specimens (3), but provides no advantage in spatial
resolution in thin specimens. This is illustrated in Fig.
5, in which the same pollen grain is imaged
with the two systems (Fig. 5, A and B). These pollen
grains are characterized by spikes or spines on their surface that
provide a convenient illustration of spatial resolution. The optical
section provided by two-photon excitation (Fig. 5B) is clearly
not as thin as that provided by the one-photon confocal system. The
lower axial resolution of the two-photon system (Fig. 2) is apparent as
increased haze around the sharp spines. We have recently published the
first confocal images of Ca2+ within individual vascular
smooth muscle cells of the walls of living pressurized resistance
arteries (11). This is a thick specimen for which two-photon imaging
should be advantageous. Accordingly, we stained a rat mesenteric artery
with an amine-reactive fluorescein dye
[5-(and-6)-carboxy-2',7'-dichlorofluorescein diacetate, succinimidyl ester, Molecular Probes, Eugene, OR], and the same region of living artery wall was imaged with the two systems (Fig. 5,
C and D). In this case, the two-photon image is
superior, revealing more clearly the individual vascular smooth muscle
cells and providing a better optical section of endothelial cells that
are folded into the arterial wall (visible as the "trough"
running down the center of the image).
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DISCUSSION |
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TOP
ABSTRACT
INTRODUCTION
METHODS AND MATERIALS
RESULTS
DISCUSSION
REFERENCES
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We present a simple, efficient, and cost-effective combined confocal and two-photon laser scanning microscope. It provides spatial resolution that is comparable to or better than that of typical commercial microscopes, while using approximately nine times less light (in confocal mode). This provides substantial benefits for physiological studies. For example, voltage-clamp studies of Ca2+ sparks in cardiac cells typically involve repetitive scanning of the same region of the cell thousands of times during a series of voltage-clamp pulses (10). The high efficiency of our custom confocal microscope means that such experimental protocols can be much more prolonged, enabling the collection of much more data, because there is less risk of photodamage to the cell and less photobleaching of the Ca2+ indicator dye. The detection of fluorescent Ca2+ indicator signals in intact arteries under physiological conditions (11) is also greatly facilitated by the efficient signal collection, because the rate of loss of fluorescent indicator dyes from smooth muscle cells is relatively high at mammalian temperatures compared with room temperature.
The use of photon counting in this microscope provides high precision
of light measurement and obviates some sources of noise (1). Most
importantly, black levels, which are very important in fluorescence
ratios (e.g., F/F0), are very low in this instrument and
are not subject to analog offsets in amplifiers. Whereas the instrument
excels at measuring low levels of fluorescence, its dynamic range is
limited by the dead times of available photon-counting systems.
Although not demonstrated here, a major advantage of this instrument over typical commercial instruments is its adaptability for physiological experiments in which various stimuli (e.g., voltage-clamp pulses) may have to be given in exact temporal relationship to image acquisition. Because all functions of the instrument are readily controlled and accessible, the timing of physiological stimuli and data acquisition is facilitated.
Finally, by incorporating both one- and two-photon fluorescence-excitation systems, versatility is obtained for studying both thin and thick specimens. The addition of a Ti:sapphire laser that is tunable to the system also makes possible the two-photon excitation of fluorophores over a wide range of wavelengths (5).
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ACKNOWLEDGEMENTS |
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We acknowledge the invaluable assistance of Dr. Victor A. Miriel with the preparation of isolated arteries used in this study.
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FOOTNOTES |
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The costs of publication of this article were defrayed in part by the payment of page charges. The article must therefore be hereby marked "advertisement" in accordance with 18 U.S.C. §1734 solely to indicate this fact.
Address for reprint requests and other correspondence: W. G. Wier, Dept. of Physiology, Univ. of Maryland School of Medicine, 655 West Baltimore St., Baltimore, MD 21201 (E-mail: gwier001{at}umaryland.edu).
Received 22 October 1999; accepted in final form 20 December 1999.
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REFERENCES |
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TOP
ABSTRACT
INTRODUCTION
METHODS AND MATERIALS
RESULTS
DISCUSSION
REFERENCES
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) or 2 APDs (
). Solid line represents an
ideal detector for which no correction would be required. B:
amplitude frequency histograms of pixel values in a 2-APD detection
system. Open and filled circles plotted in histogram represent
distributions of pixel values in each of 2 separate APDs, and the solid
line is their sum, obtained by addition in a digital adder. C:
measured intensity-response function for confocal microscope and 2-APD
detection system. 
, same data after
"correction" for system dead time. Straight lines were obtained
by least-squares fit to corrected data. D: autocorrelation
function for 2-photon excitation. "Pulse width" obtained, as
full-width at half-maximum of autocorrelation function, was
109 × 10
