28
Multi-Photon Molecular Excitation in
Laser-Scanning Microscopy
Winfried Denk, David W. Piston, and Watt W. Webb
INTRODUCTION
Multi-photon microscopy (MPM), which is based on molecular
excitation by multi-photon absorption (MPA) and is usually combined with laser-scanning microscopy (LSM), has fulfilled its early
promise (Denk et al., 1990), as evidenced by continued growth of
its application to vital imaging of biological systems (for a recent
collection of reprints, see Masters, 2003). Conventional fluorescence microscopy can provide submicron spatial resolution of
chemical dynamics within living cells, but is frequently limited
in sensitivity and spatial resolution by background due to out-offocus and scattered fluorescence. The superlinear character of
multi-photon excitation (MPE) avoids background because the
excitation is almost entirely confined to the high-intensity region
near the focal point. As excitation of the out-of-focus background
fluorescence is avoided, no confocal spatial filter is required; we
retain all of the advantages of a (single-photon) confocal microscope and gain the absence of out-of-focus photobleaching and
photodamage.
Multi-photon molecular excitation during a single quantum
event was first predicted more than 75 years ago (Goeppert-Mayer,
1931) and consists of the simultaneous absorption of multiple
photons that combine their energies to cause the transition to the
excited state of the chromophore. For example, simultaneous
absorption of two photons of red or infrared light can excite a fluorophore that normally absorbs ultraviolet (UV) or blue/green
light. The fluorophore then emits fluorescence with a wavelength
that usually is shorter than the exciting laser wavelength. Because
multi-photon absorption requires at least two photons for each
excitation, its rate depends on a higher algebraic power of the
instantaneous intensity, just as the rate of a chemical reaction,
nA + B Æ C, varies with the nth power of the concentration of A.
Because of the large intensities required, the first experimental
observation of two-photon excitation (2PE) (Kaiser and Garrett,
1961) and three-photon excitation (3PE) (Singh and Bradley,
1964) had to wait for the invention of the laser (Maiman, 1960),
more than 30 years after Maria Goeppert-Mayer’s prediction.
In the decades following Kaiser and Garrett’s work, a fair number
of spectroscopic studies using 2PE were performed (reviewed,
e.g., by Friedrich and McClain, 1980; Birge 1986), mainly to
exploit the different quantum-mechanical selection rules that
govern 2PE.
Nonlinear optical effects were first used in microscopy
to produce images of second harmonic generation in crystals
(Hellwarth and Christiansen 1974; Sheppard and Kompfner 1978).
The first MPM images (using 2PE) were reported in 1990 (Denk
et al., 1990), with the expectation from its inception to develop
nonlinear laser microscopy as a new tool for biophysical research.
Technological advances in two different areas have made
nonlinear laser microscopy, in general, and MPM, in particular,
practical: first, the development of LSM (Davidovits and Egger,
1969; Wilson and Sheppard, 1984) and, second, the development
of mode-locked lasers that are capable of generating ultrashort
pulses (ª100 fs) of red or infrared light at high repetition rates
(ª100 MHz). (For a selection of reprints on this subject, see
Gosnell and Taylor, 1991.)
Early applications of two-photon microscopy (2PM) to the
study of dynamic biochemical processes in living cells demonstrated some of the advantages for quantitative three-dimensional
(3D) and four-dimensional (4D) (space and time) resolved fluorescence microscopy (see below). Much progress has since been
made in recognizing the important fundamental parameters. Solutions have been found for technological problems such as efficient
detection in scattering samples and effective commercial instrumentation has been created. This has resulted in many important
biomedical research problems being successfully attacked using
MPM (see below), in particular those requiring visualization of
dynamic cellular processes. One of the main areas of application
has been high-resolution imaging inside highly scattering brain
tissue in vitro and in vivo.
The goal of this chapter is to elaborate on the physical principles of MPM, and to point out their relevance to actual instrument
design, including the selection of the appropriate laser light source.
We also discuss the challenges related to chromophore selection
and characterization and then list some of the applications where
MPM has made a difference.
PHYSICAL PRINCIPLES OF MULTI-PHOTON
EXCITATION AND THEIR IMPLICATIONS FOR
IMAGE FORMATION
Physics of Multi-Photon Excitation
How and why is MPE different from 1PE and how does this lead
to the unique properties of MPM? Because most aspects become
clear when considering 2PE, we will discuss mostly 2PM and point
to differences with higher orders where necessary. We will especially explore the complications involved in determining reliable
numbers for multi-photon absorption cross-sections and why and
how the temporal structure of the excitation light can affect
imaging performance.
Winfried Denk • Max-Planck Institute for Medical Research, Heidelberg, Germany
David W. Piston • Molecular Physiology Biophysics, Vanderblt University, Nashville, Tennessee 37323
Watt W. Webb • School of Applied and Engineering Physics, Cornell University, 212 Clark Hall, Ithaca, New York 14853
Handbook of Biological Confocal Microscopy, Third Edition, edited by James B. Pawley, Springer Science+Business Media, LLC, New York, 2006.
535
536
Chapter 28 • W. Denk et al.
2PE as used here refers to the simultaneous absorption of two
photons of longer, not necessarily identical, wavelengths, l1 and
l2, that combine their energies to cause a molecular excitation that
would otherwise require a single photon with a shorter wavelength
-1 -1
(l-1
1 + l2 ) . This situation is distinct from sequential two-photon
absorption (2PA), not considered here, where the molecule is
excited into an intermediate (metastable) state by the first photon,
and from there into the final state by the second photon.
The transition probability for simultaneous 2PA depends (as
mentioned above) on the square of the instantaneous light intensity. The use of brief but intense pulses, therefore, increases the
average two-photon absorption probability for a given average
incident power. It is desirable to minimize the average excitation
power to minimize undesirable 1PA, which can occur all along the
excitation beam and is usually responsible for most heating (see
below) and may also cause photodamage directly. The multiphoton “advantage” (defined below) for n-photon excitation is proportional to the inverse excitation duty cycle to the n-1 power. For
example, using 100 fs (1 fs = 10-15 s) duration pulses at a 100 MHz
repetition rate leads to 100,000-fold and 1010-fold improvements
over CW illumination for 2PA and 3PA, respectively. The use
of such short pulses and small duty cycles is, in fact, essential
to permit image acquisition within a reasonable time while using
biologically tolerable power levels. What constitutes a tolerable
power level is, however, hard to define and depends on sample
properties, as well as imaging parameters such as magnification
and scan speed. With high numerical-aperture (NA) diffractionlimited illumination, tolerable average power levels are generally
around a few milliwatts at the focal spot. Due to losses in instrument optics and sample, source laser powers of over 1 W may still
be needed for deep imaging in scattering tissue (Denk, 1996).
The probability pa that a fluorophore at the center of the focus
absorbs a photon pair during a single pulse is, using the paraxial
approximation, given by Denk and colleagues (1990):
2
2
pNA ˆ
2
x,
pa = d P Fp-1 Ê
Ë 2 phcl ¯
(1)
which depends linearly on the two-photon cross-section d, quadratically on the average power ·PÒ, on the fourth power of the
ANA, and inversely on the repetition frequency FP; l, c, h̄ are the
wavelength, the speed of light in vacuum, and the Planck quantum
of action, respectively; the two-photon “advantage” factor x, is
calculated as follows:
t2
x=
P2
=
2
P
(t1 - t2 )Ú P 2 (t )dt
t1
t2
Ê 2
ˆ
Á Ú P (t )dt ˜
Ët
¯
2
,
(2)
1
with (t1 - t2) = Fp-1. For a pulse that is Gaussian in time (see below)
with a width tp (time between the half-power points) one finds
x ª (Fptp)-10.664, and for a pulse with a hyperbolic-secant shape
the quite similar value of x ª (Fptp)-10.558.
A curious property of 2PE (but not of >2 PE) is that, in spite
of the strong NA dependence of the peak excitation rate, the total
amount of 2PE arising from a focused laser beam in a homogeneous distribution of fluorophores is independent of NA. This can
be understood intuitively by realizing that the decline of the peak
2PA probability by reducing NA is exactly compensated by an
increase in the focal volume, and thus an increase in the number
of fluorophores in the excitation region. The total absorbed power
can be calculated using a slightly modified form of Eq. 4 of Birge
(1986):
2
pabs =
dC P hx
2 phc
(3)
where C is the chromophore concentration and h is the refractive
index.
The quantum-mechanical selection rules for 2PA differ from
those for 1PA (Birge, 1979, 1986; Friedrich and McClain, 1980;
Loudon, 1983). In fact, for isolated atoms a transition allowed for
1PA would be strictly forbidden for 2PA and vice versa. However,
due to their reduced symmetry and the effect of molecular vibrations, strict parity selection rules do not usually hold for complex
dye molecules (McClain, 1971).
A number of heuristic rules for the expected two-photon
spectra can nevertheless be formulated when the single-photon
spectrum is known: (1) Some 2PE usually occurs at a particular l
whenever 1PE occurs at l/2. (2) Additional features appear, if at
all, on the short wavelength side of the spectrum. (3) Two-photon
spectra are generally broader than single-photon spectra. (4)
Good (strongly absorbing) single-photon fluorophores are often
very good two-photon fluorophores, whereas bad single-photon
absorbers tend to be very bad two-photon absorbers. The absence
of additional long-l features is simply due to the fact that, toward
longer wavelengths, the combined energy of the photons is no
longer sufficient to reach the excited state. Both rules 2 and 3 arise
because single-photon inaccessible states with higher energy that
have no direct wave-function overlap with the ground state can
often be reached with two-photon excitation through intermediate
(virtual) states that do overlap with both the initial and the final
state (Mortensen and Svendsen, 1981; Loudon, 1983; Birge, 1986).
Rule 4 arises because, as the two-photon excitation process uses
the typical transition matrix element twice, its size affects the twophoton cross-section quadratically.
Equation 1 is only correct as long as the probability Pa for each
fluorophore to be excited during a single pulse is much smaller
than one. The reason for this is that during the pulse (given a pulse
length of about 100 fs and a typical excited-state lifetime tf in the
nanosecond range), the molecule has insufficient time to relax to
the ground state, which is a prerequisite of being able to absorb
another photon pair. Therefore, whenever Pa approaches unity, saturation effects begin to occur. In a strongly focused beam with
pulse lengths and repetition rates as mentioned above, average
power levels of several tens of milliwatts were estimated to cause
saturation (Denk et al., 1990). However, the use of recently
developed fluorophores and, particularly, of quantum nanoparticles
with large two-photon cross-sections can lead to saturation at much
lower power levels. Because saturation depends on the location
within the focal spot, the point spread function is altered in a way
that reduces the resolution. For comparison, the power levels
leading to ground-state depletion in single-photon excitation are
on the order of 1 mW (see Chapter 2, this volume). Often (but not
always) the desirable time between pulses is around tf because
slower repetition rates leave the fluorophore idle between pulses,
thus lowering the saturation limit on fluorescence output, and
faster repetition rates erode the two-photon advantage x, raising
the required average power to achieve a particular fluorescence
level. Repetition rates of around 100 MHz (one pulse every 10 ns),
which are common in commercially available mode-locked lasers,
are thus in the desirable range even though somewhat higher repetition rates can reduce saturation and nonlinear bleaching effects
Multi-Photon Molecular Excitation in Laser-Scanning Microscopy • Chapter 28
when high fluorescence rates are needed and average power is not
limiting.
Saturation due to ground state depletion limits the amount of
two-photon excitation power that can be usefully directed into a
single, diffraction-limited spot and thus limits the maximally available fluorescence power and hence the signal acquisition rate.
Given sufficient laser power, simultaneous illumination of multiple focal volumes (e.g., with line or microlens-array illumination
(Chapters 10 and 29, this volume) (Brakenhoff et al., 1996;
Straub and Hell, 1998; Egner and Hell, 2000; Andresen et al.,
2001) can evade this image-rate limitation by a factor given by
the number of simultaneously illuminated focal volumes (ns).
It is, however, necessary to use either ns descanned detectors
(see below), or an imaging detector. This, in turn, precludes the
use of multi-spot excitation in strongly scattering specimens (see
below).
1
0
537
1
-200
0
200
1
1
0
0
-200
0
200
-200
0
200
-200
0
200
Optical Pulse Length
It might appear that in order to increase the two-photon advantage
the excitation pulses should be as short as possible. This is not so
mainly for two reasons: first, and of greater practical importance,
is the fact that during propagation through optical materials and
reflection off multi-layer dielectric coatings, pulses are spread in
time due to group velocity dispersion (GVD). This effect, illustrated in Figure 28.1, is due to the fact that the light in ultrashort
pulses consists of quite a range of optical frequencies, and thus
wavelengths. A 70 fs pulse centered at 800 nm, for example, is
spread over 13 nm in wavelength. For a Gaussian pulse (intensity
as a function of time t: n(t) µ exp Î- 4ln(2)(t - t0)2tp-2˚ with a pulse
width ts (between the points of half maximum intensity) the socalled “transform-limited” bandwidth (Dl), where the phases of all
wavelength components are arranged to yield the shortest pulse
possible, is related to the spread in optical frequencies (Df ) by: tp
= 2ln(2)p-1Df -1 = 0.441271/Df, where Df = cDl/(l2c), with lc the
center wavelength. GVD arises in optical materials as wave
packets of different wavelength travel with different speeds, determined by their group velocities cg = ∂w/∂k = ∂(ck/h)/k = c/h ck(∂h/∂k)/h2 (not to be confused with the phase velocity w/k = fl
= c/h) where w is the angular frequency, k and l, are, respectively,
the wave number and the wavelength inside the material, and c is
the speed of light in vacuum, and h is the refractive index. For a
given optical path, the accumulated GVD then gives rise to a
certain amount of group delay dispersion (GDD), i.e., light from
the red end of the spectrum arrives at a different, usually earlier,
time than light from the blue end. This leads to a chirped (frequency swept) pulse that is longer than the original pulse but still
contains (at unchanged spectral density) the same optical frequencies and hence wavelengths. Because the pulse’s total energy
content is unchanged by GDD, chirping always reduces the peak
intensity and hence the average squared intensity, which, in turn,
determines the two-photon excitation probability. The difference
in the arrival times increases with increasing Dl, and because
shorter pulses have a broader spectrum (see above), they are, for
a given amount of GDD, stretched more than longer pulses. This
effect is compounded by the fact that the same amount of stretching lengthens a shorter pulse by a larger fraction of its original
length t0p. Therefore, the two-photon advantage x, which depends
on tp, is degraded by a pulse-broadening factor depending on the
inverse square of t0p. For a Gaussian pulse we find for the spread
pulse
2
t p = t 0p 1 + [ 4 ln(2)l V t 2p ] ,
(4)
-200
0
200
1
1
0
0
-200
0
200
FIGURE 28.1. Simulating the effect of group velocity dispersion (GVD) on
the pulse shape of an ultrashort pulse. The pulse has initially a FWHM width
of 40 fs and is then dispersed by about 1250 fs2 of GDD (corresponding to about
35 mm of fused silica or less than 5 mm of SF59 glass). For comparison the
electric field (top row of panels), the intensity (middle row), and the squared
intensity (bottom row; corresponding to the two-photon excitation efficiency)
are shown both without (left column) and with dispersion (right column). An
unrealistically long center l (4000 nm) was chosen in order to emphasize the
chirping effect. In reality (for 900 nm light), a 30 fs pulse would be 10 full
cycles long rather than about 2 cycles as shown here.
where l is the length of the light path inside the material and V =
c-1 ∂2(hw)/∂w2 the GVD parameter. Therefore, for a given amount
of GDD (lz) there is an optimal t0p that leads to the shortest tp after
passing the group-velocity-dispersing elements. For example, for
1 cm of fused silica, with V = 362 fs2/cm at l = 800 nm, the shortest tp (ª45 fs) is obtained for t0p = 30 fs. Highly corrected lenses
often use optical glasses that have considerably larger GVDs, and
in most microscopes, light passes through considerably more than
1 cm of glass. For example, GVD values at 800 nm are 338, 453,
447, 870, 1187, 1193, 1896, 2236, and 2936 fs2/cm for the Schott
glasses FK51, BK7, BKI, LFS, SF2, TiF6, SF11, SF57, and SF59,
respectively (calculated from refractive index data in the Schott
glass catalog; Schott Glass Technologies, Duryea, PA). The GDDs
of microscope objectives and whole laser-scanning microscopes
have been explored experimentally and theoretically (Guild et al.,
1997; Muller et al., 1998; Wolleschensky et al., 1998). The effects
of GDD in MPM are discussed again below.
538
Chapter 28 • W. Denk et al.
In theory, broadening that is due to the GDD can be compensated by prechirping the pulse (giving the blue wavelengths a head
start), using a prism or grating arrangement (Fork et al., 1984) in
such a way that different wavelengths arrive at the sample almost
simultaneously after passing the microscope optics. However, in
view of the added alignment complexity and possible power losses
in the compensation optics as well as the need for readjustment to
a different GDD value for each objective lens and excitation wavelength, it has to be carefully weighed whether prechirping is worth
the effort, as it might well be if single-photon absorption or lack
of laser power are an issue. It is worth noting (Eq. 4) that for large
amount of GDD the pulse length roughly increases linearly with
the amount of GDD, but small amounts affect the pulse length disproportionately less.
Even if pulse broadening by GDD is completely compensated,
there is a second factor putting a lower bound on the optimal t0p.
As the l spectrum broadens with the shortening of the pulses, it
will eventually become wider than the absorption spectrum of the
chromophore. This limit is, however, not pressing because most
chromophores used in fluorescence microscopy have spectra
between 20 and 50 nm wide (full-width half-maximum, FWHM),
for which pulses with a length of 23 and 9 fs, respectively, have a
matched (doubled) spectral width at 700 nm. A very interesting
development in this context is the use of coherent control techniques that apply complex phase relationships between the different l components to select particular excitation pathways (see,
e.g., Walowicz et al., 2002; Lozovoy et al., 2003).
Excitation Localization
Most of the properties that make MPM so useful for fluorescence
microscopy derive from the quadratic or stronger dependence of
the excitation probability on the excitation light intensity. In a
strongly focused excitation beam, the excitation probability
outside the focal region falls off with z-2n, where z is the distance
from the focal plane and n is the number of photons absorbed per
quantum event. In a thick sample with a spatially homogeneous
distribution of chromophores and for a Gaussian beam, about 80%
of the two-photon absorption, and therefore 80% of the total fluorescence, occurs in a volume bounded by the e-2 iso-intensity
surface, which for an objective lens with an NA = 1.4 is contained
within an ellipsoid (0.3 mm in diameter and 1 mm long for l = 700
nm) or approximately 0.1 femtoliter (mm3) in volume (Sandison
and Webb, 1994). This means that MP (unlike 1P) excitation is
truly localized and as a consequence provides excellent depth discrimination, which is similar (for the 2P case nearly identical) to
that of an ideal 1P confocal microscope. Because, in contrast to
the 1P case, 3D resolution is due to the confinement of excitation
to the focal volume, out-of-focus photobleaching and photodamage and the attenuation of the excitation beam by out-of-focus
absorption do not occur, and because no spatial filter (detection
pinhole) is required, none needs to be aligned. Figure 28.2 shows
a comparison between an xz-section through a bleaching pattern
that was generated by repeated 2P scanning of a rectangular area
in a single xy-plane in a thick, rhodamine-stained Formvar layer
and an xz-section through a bleaching pattern caused by 1P scanning. In the 1P case bleaching occurs throughout the depth of the
sample.
Detection
The fact that resolution and discrimination are defined by the excitation process alone leaves substantially more freedom when
FIGURE 28.2. Confinement of photodynamic effects, such as bleaching, to
the focal slice. xz-profiles of the bleach patterns formed by repeatedly scanning
the laser focus over a single xy optical section in a thick film of rhodaminedoped Formvar until the fluorescence from the focal plane was largely
bleached. The scanned area extends through about half the image width shown.
Single-photon excitation was used on the left and two-photon excitation on the
right. 1P bleaching extends throughout the sample thickness while 2P bleaching is confined to a thin region around the focal plane. The widening of the
bleached region seen above and below the focal plan for 1P bleaching is due
to the high NA illumination cone.
choosing the detection strategy: (1) The emitted light does not have
to pass through the microscope objective at all, allowing the use
of emission wavelengths that are not transmitted by the objective
lens but could instead be detected by a photodetector placed, for
example, on the far side of the sample. (2) The emitted light does
not have to be focused. Therefore, scattering of emitted light can
be tolerated without any loss of detection efficiency or resolution
(Denk et al., 1994; Denk and Svoboda, 1997). This is especially
useful in strongly scattering samples, such as brain tissue (Denk
et al., 1994; Yuste and Denk, 1995; Svoboda et al., 1997). Furthermore, due to reduced scattering at longer wavelength, the 2P
excitation wavelength (l2ex) can be focused to an adequately defined focus (which might be impossible for the corresponding l1ex).
Only a vanishing fraction of the short wavelength (lem) and hence
strongly scattered emission light emerges unscattered (ballistic)
and could be used for confocal detection, virtually precluding the
use of 1P confocal microscopy in such samples.
(3) Non-optical signals such as photo-chemically induced
current signals in biological cells (Denk, 1994; Furuta et al., 1999;
Matsuzaki et al., 2001) or photo-induced currents from semiconductor circuits (Xu and Denk, 1997, 1999) can be used to generate optically sectioned images.
Wavelengths
The range of 350 nm < l1ex < 500 nm (700 nm < l2ex < 1000 nm)
is most widely used to excite fluorescence indicators and photoactivatable compounds. The ability of MPM to reach short UV
excitation energies beyond those reachable with 1PE (l1ex < 300
nm) has so far only rarely been used for imaging applications
(Wokosin et al., 1996b; Xu et al., 1996; Maiti et al., 1997; Williams
et al., 1999). One reason is that photodamage can occur due to MPE
Multi-Photon Molecular Excitation in Laser-Scanning Microscopy • Chapter 28
by intrinsic chromophores (Rehms and Callis, 1993) in proteins and
DNA (Williams and Callis, 1990). While photodamage has been
insufficiently studied (there are many anecdotal reports on tissue
and cell photodamage, but few non-controversial facts) it appears
that longer excitation wavelengths are better tolerated by living
cells and tissues, especially at high excitation intensities.
Resolution
Another important question is the resolution of 1PCM versus
MPM. This was first quantitatively discussed by Sheppard and Gu
(1990) and Nakamura (1993). The answer depends strongly on
whether the fluorophore or the excitation energy is held fixed.
Using the same excitation wavelengths, the 2PM, even without a
detector spatial filter (pinhole), has a slightly improved resolution
due to the lack of a Stokes-shift effect and a very small equivalent
pinhole size (Sheppard and Gu, 1994). When using, more appropriately, the same fluorophore, l2ex ª 2 ¥ l1ex , the resolution of the
2PM is degraded by a factor of almost 2 (somewhat less if the fluorophore has a large Stokes shift) compared to the ideal (zeropinhole size) confocal microscope. However, for a realistic pinhole
size (Gauderon and Sheppard, 1999), the performance of the
1PCM deteriorates, so that in practice the resolution in 1PCM and
MPM is about the same. A significant resolution enhancement in
multi-photon microscopy, albeit at the expense of collection efficiency (see below), can be achieved by using a detection spatial
filter in conjunction with a relatively short excitation wavelength
(Stelzer et al., 1994). The resolution can be substantially improved
in both 1PM and MPM along the axial direction by using illumination from almost all directions as with the 4-Pi microscope (Hell
and Stelzer, 1992) (see also Chapter 30, this volume).
In conclusion, 2PE does not normally lead to resolution
improvements over confocal microscopy. In fact, if resolution is
of paramount importance and scattering is moderate, 1PE confocal microcopy is usually better. Recently it has been shown that
stimulated emission depletion (STED) microscopy, a very different type of nonlinear optical microscopy, can overcome the farfield diffraction limit (Dyba and Hell, 2002) (see also Chapter 31,
this volume).
Photodamage: Heating and Bleaching
Photodamage to cells and tissue can result from 1PA or MPA,
depending on illumination wavelength, on type and concentration
of chromophores present, and on the power level. In particular,
when infrared ls are used (>900 nm), we have to consider heating
due to increased absorption by liquid water, which is not a problem
at visible and near-UV ls where water is very transparent
(e.g., see Fig. 3 in Svoboda and Block, 1994 or Fig. 23.3, this
volume).
For 1PE, an upper-bound estimate of the temperature rise can
be made using a 2D approximation because absorption occurs all
along the beam path. The calculation sketched here is the same as
was used to analyze thermal lens effects (Whinnery, 1974; Kliger,
1983, and references therein). For the temperature rise T at the
center of the beam (r = 0) as a function of the time t after switching on the beam one obtains
T2D (t , r = 0) =
aP
Ê 2t
ˆ
lnÁ + 1˜
¯
4 pkT Ë t c
(5)
where a is the absorption coefficient, P the laser power, kT the
thermal conductivity, and tc = w 20/(4k) the thermal time constant,
which is a measure of how fast steady-state conditions are ap-
539
proached and which depends on w0, the Gaussian beam parameter
and is equivalent to the beam radius (1/ez intensity) in the focal
plane, and on the thermal diffusivity k = kT/r where r is the volume
heat capacity. For a diffraction-limited beam at high-NA (wo = 200
nm), tc ª 70 ns in water (using kw = 0.6 WK-1 m-1, kw = 1.44 ¥ 10-7
m2s-1) and for absorption by pure water, the pre-factor in Eq. 4
is 0.013, 0.21, and 0.66 K at ls of 700, 1000, and 1300 nm,
using the absorption coefficients for water of 0.02, 0.32, and
1.0 cm-1, respectively; the laser power was assumed to be 50 mW
(approximately the saturating intensity; Denk et al., 1990). Slightly
lower values, still logarithmically diverging with time, are found
if axial heat transport is taken into account (Schönle and Hell,
1998). Due to the small beam diameter, tc is rather short (ª70 ns)
for high-NA objectives. For video-rate scanning microscopes
(Goldstein et al., 1992; Fan et al., 1999; see Chapter 29, this
volume) this results in a temperature rise of only 1.55 times the
pre-factor but at 10 ms dwell-time (typical for non-resonant mirrorscanned instruments), the temperature rise is 5 times the prefactor.
For an infinite sample no steady-state value for the temperature
would ever be reached. In practice, the temperature rise will eventually be limited, by the finite sample size and by convection or
bath perfusion, which remove heat at a rate much faster than heat
conduction alone. For stationary applications or when continuously scanning a small area, rather large logarithmic factors can
occur, however. Therefore, water absorption may have to be taken
seriously, particularly at high illumination powers and long wavelengths or when attempting multi-photon excitation with CW
lasers (Hanninen et al., 1994, 1996; Booth and Hell, 1998; Hell et
al., 1998). Fast scan rates, rapid bath perfusion, thin sample cells,
and, of course, maximizing the two-photon advantage using the
shortest pulses possible are remedies to reduce high peak temperatures. To assess the 1P effects of infrared (IR) beams on biological specimens, we can also exploit the experience gained with
optical tweezers, which are routinely used on living cells
at comparable or higher power densities (Ashkin et al., 1987;
Svoboda and Block, 1994) and for which damage has been
assessed for most of the wavelength regime used in 2PM (Neuman
et al., 1999).
Heating due to 2PA is restricted to the focal region. A 3D model
is, therefore, appropriate. Because we are interested in the case of
high-NA, we can use the approximation that the release of heat
occurs uniformly within a sphere with radius w0 centered at the
focus. The relationship one gets for the temperature rise is:
2 tc ˘
È
(6)
ÍÎ1 - 2 t + 3t ˙˚
c
where Pabs is the total absorbed power (see below). For large t,
when the square root goes to zero, T3D, unlike T2D, approaches
an asymptotic value, given by the factor in front of the square
brackets.
For high energy (mJ) pulses at low repetition rate, the local
temperature rise during a single pulse can easily be large enough
to cause damage, but we know little about how damage might be
exacerbated for the case of pulsed light at high repetition rates
(fR ª100 MHz) compared to the CW case with the same average
power. Because tc is longer than the interpulse interval (1/fR), the
incremental temperature rise during a single pulse is smaller than
the steady-state temperature rise T3D (t = •) roughly by a factor of
tcFr.
We conclude that heating during high-repetition-rate pulsed
illumination can largely be treated like CW illumination and is
often negligible at practical 2PM parameters. Attention has to be
paid to situations where high local concentrations of chromophore
T3D =
Pabs
w 0 4 pkT
540
Chapter 28 • W. Denk et al.
occur, as, for example, for DNA stains, which can bind at a concentration of one per base pair or where equilibration of the molecular temperature with its environment cannot be automatically
assumed (Akaaboune et al., 2002).
Because of the localization of excitation to the focal volume,
total photobleaching in MPM is generally much reduced compared
to 1PE microscopy. However, it has been shown that an increased
photobleaching rate from within the focal volume can occur by
a mechanism where the fluorophore is initially excited by simultaneous 2PA, and then one or more photons interact with the
excited molecule, possibly via higher-order resonance absorption
(Patterson and Piston, 2000). This effect can be quite pronounced
for readily photobleachable dyes, such as fluorescein, where the
difference between one- and two-photon photobleaching rates can
be a factor of 10 at the power levels that are typically used in
biological imaging (100 mW CW for single-photon excitation, and
3 mW 150 fs pulses for MPM). However, for more stable dyes,
such as the green fluorescent proteins (GFPs), carbocyanines, and
AlexaFluors, the photobleaching rate is in our experience often too
small for the difference to be measurable at the usual imaging
intensities. While in some cases direct higher-order absorption
(three or more photons) may be relevant, several studies (Koester
et al., 1999; Konig et al., 1999) have found that longer pulses
(which reduce higher-order absorption) do not reduce the damage
done per excitation event.
INSTRUMENTATION
options for detection. In this section we will discuss laser sources
suitable for MPE, the advantages and drawbacks of the various
methods of detecting fluorescence and other contrast signals, and
specific problems that occur with non-mechanical (e.g., acoustooptical) beam power control and deflection. We assume that the
reader of this chapter is familiar with the principles of 1PCM (other
chapters, this volume). Short shrift will, therefore, be given to those
aspects such as mechanical beam scanning, data collection,
storage, and display that are largely identical for 1PCM and MPM
instruments. The potential user should also be aware that MPMs
are relatively easy to set up and are now available as integrated
systems from several manufacturers. MPM systems are still expensive with the price of the laser system (>$150,000) being between
one third to one half of the total system cost. With a mode-locked
laser, one has, however, also acquired the light source necessary
to do time-resolved fluorescence measurements (Piston et al.,
1992; Zhang et al., 2002; and Chapter 27, this volume).
Lasers and the Choice of Excitation Wavelengths
CPM Laser
The first 2P images (Denk et al., 1990) and 2P photochemical
microcopy images (Denk, 1994) were recorded using collidingpulse mode-locked (CPM) lasers (Valdmanis and Fork, 1986) at
615 nm excitation wavelength. Today this laser type is of only historical interest.
Hybrid Mode-Locked Dye Laser
Setup (Fig. 28.3) and operation of a MPM system are very similar
to those of a 1P laser-scanning microscope. The main differences
lie in the type of excitation lasers and in the increased number of
Another early type of ultrashort pulse dye laser system is the
hybrid mode-locked dye laser. These systems use an actively
mode-locked argon-ion or a frequency-doubled neodymium: YAG
dichroic
mirror
scan
mirrors
mode-locked
laser
eyepiece
filter
pinhole
time scales
pulse repeat
PMT
descanned
detection
-8
10 s
PMT
dichroic
mirror
objective lens
10-9s
10-13s fluorescence
pulseemission
width
non-optical
detection
transfer
lens
PMT
whole-area
detection
external
detection
FIGURE 28.3. Schematic diagram of a two-photon laser-scanning microscope illustrating various detection possibilities. The stream of incoming laser light
pulses is raster scanned (xy scanner, only one axis is shown here) in a way that is identical to the single-photon LSM. For fluorescence microscopy, several detection possibilities are indicated: (1) external: fluorescence light bypasses objective lens; (2) whole-area: fluorescence light passes objective lens and is then
deflected by a dichroic mirror to be focused onto the detector by a transfer lens; (3) descanned: as in the 1PLSM, the fluorescence light is reflected off the scanning mirrors, allowing confocal detection (see text). Not shown, but possible and occasionally used, is focal-array detection, where, after deflection by a dichroic
mirror, fluorescence light is detected by an array detector located in an image plane. Yet another possibility is non-optical detection using, for example, an electrically recorded signal from the sample. Time scales are indicated in the left inset.
Multi-Photon Molecular Excitation in Laser-Scanning Microscopy • Chapter 28
laser to pump a dye laser that also contains an intracavity saturable
absorber jet. Such systems are rather expensive and difficult to
operate and are therefore rarely used for multi-photon imaging.
The remaining advantage over the titanium : sapphire laser (see
below) is the access to the range 550 nm < l < 700 nm, which is
desirable for some uncaging experiments but has been virtually
abandoned for imaging due to photodamage problems (Kiskin
et al., 2002).
Titanium : Sapphire Laser
For most applications, the light source of choice for MPM currently
is the self-mode–locked titanium : sapphire (Ti : Sa) laser (Spence
et al., 1991), nowadays pumped by a frequency-doubled diodepumped Nd : Vanadate laser rather than a power- and cooling-water
hungry argon-ion laser. The Ti : Sa laser provides a large tuning
range (from slightly below 700 nm to slightly above 1050 nm) with
pulse lengths shorter than 100 fs and sufficient power (2 W average
at the peak of the tuning curve, down to a few hundreds of milliwatts at the edges when pumped with 10 W) to permit saturating
excitation (see Physical Principles) of most fluorophores with a
high-NA objective over much of the laser’s tuning range. The
tuning range of Ti : Sa is now covered by a single set of cavity
mirrors, with optics changes only required to reach wavelength
above 1000 nm or below 700 nm. Currently, several manufacturers
offer turnkey laser systems that contain the pump source and Ti :
Sa laser inside a single housing, are computer controlled, and no
longer require any mechanical adjustments by the operator.
Other Light Sources
If losses in the excitation path are too large, it is sometimes not
possible to achieve the desired excitation rates with the multiphoton advantage factors available for a laser oscillator alone. A
reduction in repetition rate while maintaining average power can
then increase the excitation efficiency substantially (Beaurepaire
et al., 2001). This can be achieved by increasing the cavity length,
cavity dumping, or regenerative amplification. The last approach
has recently been shown to allow imaging down to the surfacegenerated-background limit (Theer et al., 2003). Direct use or frequency doubling of femtosecond pulses from optical parametric
oscillators (OPO) (Cheung and Liu, 1991; Fu et al., 1992; Powers
et al., 1994; Keller, 1996) may provide an almost universal, if
expensive, solution to cover almost all of the desired wavelength
range.
One factor limiting multi-photon microscopy is the cost of the
laser source, which, in spite of early hopes, has not come down
significantly with the introduction of diode pumping (for a review,
see Keller, 1994). One reason is that gain materials that can be
directly diode-pumped (Keller, 1996) have insufficient tuning
ranges and/or unfavorable thermal characteristics.
In niche applications, other sources (Wokosin et al., 1996a)
have been used, partly within, for example, the Cr : LiSaF laser
(Svoboda et al., 1996a) or outside, for example the Cr : Forsterite
laser (Liu et al., 2001), the Ti : Sa tuning range.
Excitation Wavelengths
One reason for the success of the Ti : Sa laser for MPM is that the
range 700 nm < l2ex < 1050 nm (corresponding to 350 nm < l1ex <
525 nm) covers the range of excited state energies for many commonly used fluorophores (see below). Much shorter wavelengths,
in particular l2ex < 640 nm, are likely to cause photodamage due to
intrinsic absorption, for example, by tryptophan rich proteins
(Rehms and Callis, 1993). To minimize scattering one might
541
lengthen the excitation wavelengths and take advantage of a dip
in the absorption spectrum of water around 1040 nm, which is well
known from optical trapping experiments (Svoboda and Block,
1994). Fortunately, a large selection of microscope lenses has
become available with excellent transmission and optical correction in the near IR (Chapter 7, this volume). The use of older lenses
that were not designed for the infrared can be problematic, particularly in the case of highly corrected lenses (Neuman et al., 1999),
where non-optimal performance of the antireflective multi-layer
coatings on each of the numerous internal surfaces can reduce
overall transmission catastrophically.
Beam Delivery and Power Requirements
In general, the laser is mounted on the same vibration isolation
platform as the microscope because delivery of ultrashort pulses
through standard, single-mode fibers, which is possible in principle (Wolleschensky et al., 1998) (see also Chapter 26, this volume),
requires substantial technical efforts to prevent unacceptable
pulse broadening at the laser powers routinely required. Development of special optical fibers, such as photonic band gap fibers or
large cross-section single-mode fibers (Helmchen et al., 2002;
Ouzounov et al., 2002), may facilitate MPM applications where
fiber delivery is essential (Helmchen et al., 2001). In non-absorbing, non-scattering samples saturating pulse energies (corresponding to several tens of milliwatts of average power) can easily be
reached over most of the Ti : Sa tuning range even with 5 W of
green pump power. However, one rarely has more than the desirable power in scattering samples such as in brain slices or the intact
brain. Power availability may also be limiting when attempting to
optimize resolution by overfilling the objective back aperture.
Detection
As discussed above, excellent 3D localization is accomplished
in MPM by excitation alone. This allows more flexibility in the
optical design and, as a consequence, considerable improvements
of fluorescence collection efficiency are possible compared to the
1PCM. Figure 28.3 depicts the various options. The positions of
non-imaging detectors are designated PMT because photomultiplier tubes are usually the detectors of choice for MPM. In general,
considerations as to which detector type to use in MPM are quite
similar to those for 1PCM, and the reader is referred to Chapter
12 of this book. Among non-imaging schemes (one or a few detector elements), the main distinction is whether the emitted light
passes back through the scanning mirrors (descanned detection) or
whether the detector is sensitive to emitted light from the whole
image area at all times (whole-area detection). A variant of the
latter is external detection, where detected light does not pass
through the objective lens.
Whole-Area and External Detection
Whole-area detection (WAD) (Piston et al., 1992, 1994) is now the
detection mode of choice in the majority of MPM applications.
The WAD pathway uses a dichroic mirror somewhere between the
scanner and the objective to separate excitation and fluorescence
(alternatively the excitation light can be coupled in by reflection
from a dichroic), preferably after a minimum number of optical
surfaces to maximize detection efficiency. The signal is then
passed through the collection optics, which needs to avoid
vignetting. If the back aperture of the objective is conjugate to the
photocathode of the PMT the effect of spatial heterogeneities in
the photocathode sensitivity is reduced. One of the main advantages of WAD is the ability to efficiently collect fluorescence from
542
Chapter 28 • W. Denk et al.
specimens that scatter light at lem so strongly that only a very small
fraction can be refocused for confocal detection (Denk et al., 1994;
Denk and Svoboda, 1997; Beaurepaire and Mertz, 2002).
WAD is as vulnerable to contamination from ambient room
light as is widefield imaging with highly sensitive cameras. One
thus loses a convenient but rarely essential advantage of confocal
imaging. While WAD through the excitation lens is usually the
most convenient and efficient mode, external detection, where the
detected light bypasses the objective lens, can be necessary when
light needs to be detected that cannot (e.g., because it is of too
short a wavelength) or did not (e.g., because it went off in the
wrong direction) pass the objective. Combining through-the-lens
collecting with collecting the light passing through the condenser
has been used successfully to increase the signal-to-noise ratio in
embryo (Denk et al., 1997) and in brain slice imaging (Koester
and Sakmann, 1998; Mainen et al., 1999a, 1999b).
Another disadvantage of WAD is that detectors with a large
“phase-space volume” (given by the product of detector area and
acceptance angle) are needed thus ruling out the use of small-area
photon-counting avalanche photodiodes (Tan et al., 1999) or of
spectrometers (Lansford et al., 2001).
Descanned Detection
When converting a confocal microscope to multi-photon operation
(Denk et al., 1990) descanned detection naturally results. While
this mode is less efficient than WAD, even for clear specimens,
descanned detection does allow the use of detectors with small
apertures such as avalanche photodiodes or spectrometers. A
pinhole that is several times larger than the optimal confocal size
can be useful for excluding room light contamination from the
detected signal while still being near optimal for signal collection.
The use of a confocal pinhole as a tight spatial filter in addition to multi-photon excitation (Stelzer et al., 1994) is rarely used
because it is fraught with several drawbacks: (1) A pinhole small
enough to produce any substantial increase in resolution causes
a large drop in detection efficiency due to the fact that the
diffraction-limited volume at the l em is smaller than the excitation
volume determined by the l2ex because l em < l2ex. Such a loss of
detection efficiency is particularly serious because fluorescence
imaging of living specimens is often limited by photobleaching
and photodamage. A technically complex yet feasible solution to
this problem might be to use a small array of detectors together
with the appropriate deconvolution algorithms (Sheppard and
Cogswell, 1990).
Chromatic aberration, already a problem in 1PCM, is exacerbated in the confocal operation of MPM because the typical shift
between l2ex and lem is much larger (50 nm < lem - l1ex < 200 for
1PE, 200 nm < l2ex - lem < 500 in 2PE, and further increasing
with >2 PE).
Non-Optical Detection
A number of non-optical detection schemes have become very
promising owing to the high degree of spatial localization achieved
during excitation alone. Two-photon scanning photochemical
microscopy (Denk, 1994; Furuta et al., 1999; Matsuzaki et al.,
2001) generates images of receptor distributions by locally releasing agonists such as neurotransmitters from “caged” precursors
and detecting the agonist-induced ionic current in voltage-clamped
cells. In fact this concept was one of the motivating factors for the
initial development of MPM. Opto-acoustic detection, which has
been used to measure two-photon absorption coefficients (Patel
and Tam, 1981; Bindhu et al., 1998) could be used to measure
spatially resolved absorption that is not accompanied by fluorescence or induced current, but has to date not been tried as a contrast mechanism in MPM.
Focal-Plane Array Detection
A rather different strategy, which does not rely on scansynchronized detection to build up the image, is the use of an
imaging detector. As in conventional fluorescence microscopy, the
fluorescence is refocused to an image plane, and the image is
generated by spatially sorting the fluorescence photons into the
pixels of an array detector such as a charge-coupled device (CCD).
The lateral resolution is then determined solely by lem, which is
considerably shorter than l2ex. The optical sectioning effect due to
two-photon excitation is, however, retained and provides discrimination and resolution in z-direction. This method is the equivalent
of widefield fluorescence microscopy with only a thin focal slice
rather than the whole thickness of the sample excited. Focal array
detection is particularly useful in connection with multi-point illumination, where it allows the acceleration of image acquisition
(Straub and Hell, 1998; Egner and Hell, 2000; Andresen et al.,
2001; Fittinghoff et al., 2001; Hell and Andresen, 2001; Nielsen
et al., 2001; Egner et al., 2002). The main disadvantage of focal
array detection is that, different from the case of single-point
scanning, with whole-area detection, scattering of fluorescence
light leads to an immediate degradation of image contrast and
resolution.
Optical Aberrations
Aberrations inherent in the microscope and spherical aberration
introduced by focusing through refractile layers such as the coverslip, immersion oil, and sample (Sheppard and Cogswell, 1991;
Hell et al., 1993) broaden the focus, shift the apparent focal point
(Visser et al., 1991), and reduce the peak excitation intensity. Due
to the mathematical equivalence between the optical transfer function of the non-confocal 2PM and that of the confocal 1PM
(Sheppard and Gu, 1990; for a minor modification, see Visser et
al., 1991), the effects of monochromatic aberrations, such as
spherical aberration and astigmatism, on the amplitude and resolution of the detected signal are the same in both cases. In the
two-photon case, however, the number of molecular excitations is
actually reduced due to the smeared-out focus spot and the intensity-squared dependence of the excitation probability (see above).
When photobleaching or photodamage are the limiting factors, this
can provide a significant advantage of MPM over the 1PCM case,
where the same number of excitations occur, but the fraction of
the emitted light that reaches the detector is reduced. Nonetheless,
one must take the same precautions with MPM as with 1PCM
when interpreting absolute light levels as a function of focusing
depth. Note also, that most aberrations become rapidly less severe
as NA is reduced. The best way to avoid spherical aberration in
aqueous specimens, even at high-NA, is the use of water-immersion objective lenses, which are now widely available corrected
even for the IR range (Chapter 7, this volume).
A significant motivation for the development of 2PM was the
circumvention of the poor chromatic correction then found for
most microscope lenses in the UV. Chromatic aberration problems
play a role in (non-confocal) 2PM only in connection with the
broad l spectrum of ultrashort pulses (see above). However, this
spread is generally smaller than a typical Stokes shift and chromatic correction is easier in the IR where glass dispersion flattens
out.
Multi-Photon Molecular Excitation in Laser-Scanning Microscopy • Chapter 28
Pulse Spreading Due to Group Delay Dispersion
As discussed above, the optical materials comprising the microscope optics cause the excitation pulses to spread in time and thus
become less efficient in exciting multi-photon transitions. The
group delay dispersion (GDD) has been measured for some
objectives (Guild et al., 1997; Squier et al., 1998). As mentioned
above, the optical effort needed to generate the prechirping necessary to compensate for the GDD has to be weighed against the
improvements expected. As a general rule, GDD compensation
will be helpful or even essential when laser power is limiting, such
as for deep tissue imaging, or when single-photon absorption contributes to damage. If coherent control techniques are used, complete dispersion control is, of course, essential but then the optics
used to tailor phases can be employed for dispersion control as
well.
Control of Laser Power
For slow control of the laser power, mechanically actuated devices
such as filter wheels, graded neutral density filters, or rotating halfwave-plate/polarizer combinations (Denk, 2001) can be used.
Faster shuttering (e.g., in order to blank the beam during retrace)
or modulation requires non-mechanical devices such as acoustooptical (AO) or electro-optical (EO) modulators (i.e., Pockels
cells), which can respond on the microsecond and even nanosecond timescale (Chapter 3, this volume). EOMs achieve high
throughput but often incomplete extinction, while AOMs are
lossy, and due to limited diffraction efficiency, their extinction is
excellent.
A few problems arise specifically when ultrashort pulses are
used together with such devices: (1) In AOM, AOD, or acoustooptic tuning filter (AOTF) devices, an acoustic wave diffracts the
incoming beam by an angle that is dependent on lex. For ultrashort
pulses, which are spectrally broad, the focus, therefore, becomes
distorted and diffraction efficiency is reduced. (2) Both EOMs and
AOMs use dispersive materials, which spread the laser pulse
temporally (see above).
While the temporal spread can be easily compensated (in a few
cases multi-photon microscopy setups already contain GDD compensators), it is much more difficult but not impossible to compensate for the angular spread in an AOM (Lechleiter et al., 2002).
Limited extinction from the EOMs is often not a serious problem
because the quadratic intensity dependence of two-photon excitation allows even a moderate power reduction ratio to translate into
almost complete elimination of unwanted excitation.
Resonance and Non-Mechanical Scanning
The time resolution of closed loop galvanometer scanners is sufficient for most applications, in particular if a limited number of measurement points can be selected. However, because the time per
line cannot be reduced significantly below the about 1 ms with
closed-loop scanners, scan times for large areas can become too
long for the time resolution desired. One solution to this problem
is the use of resonant galvanometer scanner (Fan et al., 1999; see
also Chapters 3 and 29, this volume) which provide a fixed line rate
about 10 times faster, albeit at some loss of flexibility. Acoustooptical scanning (Art and Goodman, 1993) requires correction of
the diffractive spread of the wavelengths comprising short-pulse
light (Lechleiter et al., 2002), but has the advantage of more rapid
access (still limited by the acoustic transit time across the diffraction medium) and allows both scanning and intensity control.
543
CHROMOPHORES (FLUOROPHORES AND
CAGED COMPOUNDS)
The criteria for choosing, or designing, fluorophores for MPM are
essentially the same as for any other fluorescence microscopy technique: large absorption cross-section at convenient lexs, high
quantum yield, low rate of photobleaching, and minimal chemical
or photochemical toxicity to living cells. In the early days of MPM,
a heuristic approach prevailed and fluorophores were selected that
had proven useful in widefield fluorescence microscopy or 1PCM.
In most cases, two-photon excitation was found whenever there is
single-photon absorption at a l corresponding to twice the energy
of the excitation photons. Most MPM imaging still uses conventional fluorophores, and we now have two-photon spectra for many
of these (Xu and Webb, 1996; Xu et al., 1996; Zipfel et al., 2003).
On the other hand, there is a considerable effort to generate chromophores tailored to MP excitation using a donor–acceptor–donor
or acceptor–donor–acceptor strategy. These molecules maximize
the electrical dipole transition by electron transfer over relatively
long distances from donor to acceptor. By this approach, molecules
can be created with two-photon excitation cross-sections about 10fold greater than conventional fluorophores (Albota et al., 1998b;
Ventelon et al., 1999). Nanoparticles, also called quantum dots
(Bruchez et al., 1998; Han et al., 2001), which offer broad excitation spectra, but very narrow emission spectra, have the largest
measured two-photon cross-sections seen to date. This allows their
detection at very low concentrations, even in vivo (Larson et al.,
2003).
Another notable development is the movement to longer wavelengths. While in the early days of MPM the emphasis was on UVexcited dyes that were 2P-excited by red lasers, the emphasis now
is on fluorophores normally excited by visible light and 2P-excited
by IR light. This trend is mainly driven by the desire for lower
background fluorescence and deeper penetration into scattering
tissue.
Two-Photon Absorption Cross-Sections
Differences between one- and the two-photon excitation spectra
have been exploited in molecular spectroscopy because they
provide additional information about the structure of excited states.
These differences can be quite significant, see, for example, the
case of Bis-MSB (Kennedy and Lytle, 1986) or the aromatic amino
acids tyrosine and phenylalanine (Rehms and Callis, 1993), but
note the spectral similarities for tryptophan. As a rule of thumb, in
symmetrical molecules one expects l2ex < 2lex.
Calculations of two-photon cross-sections are difficult to
perform for complex molecules. Direct experimental measurements of multi-photon absorption are equally difficult because
even under optimal conditions, the fraction of the incident power
that is absorbed is rather small (using Eq. 4 we find, e.g., pabs/p =
3 ¥ 10-5 for a chromophore with a cross-section of 10-50 m4s photon-1, at a concentration of 10 mM and a laser power of 100 mW
with a two-photon advantage of 105). While thermal lensing or
acousto-optical techniques have been used to measure two-photon
absorption (Kliger, 1983), these techniques are much more complicated than single-photon spectrophotometry.
For fluorescent molecules, the shape of the two-photon excitation spectrum can be determined by detecting the intensity of fluorescence emission as a function of excitation wavelength. In order
to determine the action spectrum, the incident average laser power
(Pi), the probability of detecting fluorescence photons, and the two-
544
Chapter 28 • W. Denk et al.
photon advantage x (Eq. 2) need to be known (Xu et al., 1995).
The absolute value of the two-photon absorption cross-section can
then be calculated using the fluorescence quantum yield. Quite a
number of measured spectra are now available in the literature
(Xu et al., 1996; Albota et al., 1998b also includes URL.)
While precise calculations of two-photon absorption crosssections are difficult, several new fluorophores with particularly
large two-photon absorption cross-sections have been designed
using theoretical considerations (He et al., 1995, 1997; Marder et
al., 1997; Albota et al., 1998a; Ventelon et al., 1999, 2002;
Adronov et al., 2000; Kim et al., 2000; Zojer et al., 2002). Before
such fluorophores can come into common use, however, problems
with water solubility, derivatization, etc., will have to be solved.
For the fluorophores studied so far, the spectra of the emitted
fluorescence were found to be essentially independent of whether
excitation occurs via single- or two-photon excitation (Curley
et al., 1992). This is not surprising because the molecular relaxation process (on the picosecond scale) almost always occurs to
the same state (the lowest excited singlet state) prior to the emission (on the nanosecond scale) and therefore erases the memory
of the excitation pathway and energy.
Caged Compounds
Two-photon absorption spectra for caged compounds are more difficult to measure than those for fluorophores because the amount of
uncaged material generated is too small to be easily measured with
most analytical techniques. In some cases, uncaging can be detected
when fluorescence assays for the released agonist exist, such as for
caged ATP (Denk et al., 1990), when the product itself is fluorescent, as it is with caged fluorescein (Svoboda et al., 1996b), or when
biological effects can be detected, such as the opening of ion channels by the two-photon–induced release of caged neurotransmitters
(Denk, 1994; Matsuzaki et al., 2001; Kasai et al., 2002). Photochemical reactions are often much slower than fluorescence
emission and their speed can strongly depend on the chemical environment such as pH and ionic strength (Milburn et al., 1989; Corrie
and Trentham, 1993; Kao and Adams, 1993). The speed of release
is important for at least two reasons: (1) The pixel dwell-time must
be at least as long as the duration of the signal used to generate image
contrast, which at best is as fast as the photochemical reaction rate;
(2) diffusion of the released agonist tends to blur the image and thus
prevents high-resolution mapping. A delay of 10 ms, for example,
allows the released agonist, typically a small organic molecule with
a diffusion constant of 5 ¥ 10-9 m2 s-1, to diffuse a distance of about
3 mm (Kiskin et al., 2002).
CELL VIABILITY DURING IMAGING
The survival of the biological sample while it is being imaged is
one of the most important constraints on the usefulness of any vital
microscopy technique. While one of the reasons for pursuing MPM
as a new technique was the expectation of greatly reduced photodamage (Denk et al., 1990), it has to be kept in mind that in the
focal plane, for a given excitation rate the damage is expected to
be at least as large for 2P as it is for 1P excitation. This is because
any effect due to reactions initiated from the excited state of the
chromophore are independent of the mode of excitation. Furthermore, it cannot be ruled out that some endogenous biological molecules have unusually large two-photon cross-sections (such as
bacteriorhodopsin; Birge and Zhang, 1990) and are, therefore, particularly susceptible to damage. Another concern is the possibility
of excited state absorption, particularly at excitation rates near
saturation.
Considerable work has been performed in this area since the
first edition of this book. Two-photon excitation, particularly when
using wavelengths below 800 nm (Konig et al., 1996; Oehring
et al., 2000) (see Chapter 38, this volume) can, not surprisingly,
generate reactive oxygen species, which are implicated frequently
in photodamage (Tirlapur et al., 2001). On the other hand, when
using longer wavelengths (1064 nm), generation of reactive oxygen
species by flavin-containing proteins seems to be greatly reduced
compared to single-photon excitation (Hockberger et al., 1999).
At higher excitation levels, a steeper than quadratic power
dependence is often found both for cellular photodamage (Koester
et al., 1999; Oehring et al., 2000; Hopt and Neher, 2001) and for
photobleaching (Eggeling et al., 1998; Patterson and Piston, 2000).
It appears, however, that the damage nonlinearity is not instantaneous (i.e., three- or four-photon excitation) because for the same
mean two-photon excitation rate no change in the damage is seen
with pulse width (Koester et al., 1999; Konig et al., 1999).
There is virtually no experimental indication that heating by
water absorption (discussed in Physical Principles) is a limiting
factor in multi-photon microscopy. Heating may yet become an
issue as substantially longer wavelengths are beginning to be used
for the excitation of long wavelength fluorophores. A number of
explicit examples show an actual and significant reduction of photodamage when using two-photon rather than single-photon
imaging in biological specimens such as cultured cells
(Hockberger et al., 1999), cardiac myocytes (Niggli et al., 1994b;
Piston et al., 1994), and mammalian (Squirrell et al., 1999) and
invertebrate embryos (Summers et al., 1993).
The experience of many a microscopist is that live-cell
imaging can often be performed by reducing the excitation light
intensity to the lowest possible level, using efficient optics and sensitive detectors (Chapters 17, 19, and 29, this volume). The experience in 2PM is similar, but the range of imageable specimens is
larger. For example, in both the sea urchin (Piston et al., 1993) and
hamster embryos (Squirrell et al., 1999), two-photon excitation
allows extended observation of embryonic development, under
conditions where single-photon excitation is unsuccessful. In
another case, as part of a direct comparison of scanned laser UV
and two-photon excitation (Niggli et al., 1994a; Piston et al.,
1994), it was found that two-photon excitation allowed imaging of
the calcium indicator dye Indo-1 continuously for 5 min without
compromising cell viability. Equivalent single-photon scanning
with UV light resulted in considerable photobleaching, and over
80% cell death (Piston et al., 1994).
Those studies indicate that, even though damage is less than
with conventional UV illumination, cultured-animal-cell viability
can be compromised by two-photon excitation. Particularly worrying, and as yet unresolved, is the observation that at high illumination levels the two-photon photobleaching rate can increase
much faster than the excitation rate (Patterson and Piston, 2000),
even though it is not known whether there is a corresponding
increase in phototoxicity and whether these highly nonlinear
bleaching phenomena are limited to certain narrow classes of dyes,
such as the xanthene dyes.
A question that often arises is how to determine the mechanism
of damage. Important information is provided by its power dependence (Neuman et al., 1999; Hopt and Neher, 2001). For example,
two-photon photochemical damage should be proportional to the
square of the incident power. While a linear power dependence all
but rules out two-photon effects, a superlinear dependence on the
average excitation power could result from single-photon absorp-
Multi-Photon Molecular Excitation in Laser-Scanning Microscopy • Chapter 28
tion coupled with a nonlinear mediator for damage. Thermally
induced damage can have a rather sharp temperature threshold due
to cooperative phenomena such as protein denaturation. A definitive distinction between single- and multi-photon absorption is their
dependence on pulse length; if the pulse length is varied by introducing a variable degree of GDD (see above), the spectrum, and
hence the amount of linear (single-photon) absorption, remains
completely unchanged while 2PA drops.
Knowing the mechanism of damage is, of course, crucial for
choosing the optimal excitation strategy. For example, to reduce
single-photon, dose-rate–independent damage, a reduction of Fp
might seem appropriate in order to increase the two-photon advantage but the peak temperature during each pulse increases as Fp-1,
and can become larger than the thermal time constant. Unpleasant
surprises could also arise from additional absorption by molecules
already in the excited state (something that is more likely to occur
when operating closer to saturation) or from proximity effects
mediated by free radicals (Konig et al., 1996; Hockberger et al.,
1999; Koester et al., 1999; Konig et al., 1999; Oehring et al., 2000;
Hopt and Neher, 2001; Tirlapur et al., 2001).
APPLICATIONS
MPM has been used to address questions in quite a few areas of
biology. Particularly the imaging of intact tissue has benefited from
the properties of the multi-photon (predominantly two-photon)
microscope.
Calcium Imaging
Intracellular messenger dynamics, such as calcium ion concentration has been measured in single cells (Piston et al., 1994), but the
particular advantages of MPM over single-photon techniques
come to bear most in scattering tissue such as brain slices (Denk
et al., 1995, 1996; Yuste and Denk, 1995; Mainen et al., 1999b;
Sabatini and Svoboda, 2000; Wang et al., 2000; Oertner et al.,
2002), the stomatogastric ganglion (Kloppenburg et al., 2000), and
in vivo (Svoboda et al., 1997, 1999; Debarbieux et al., 2003). In
isolated retina 2PM allowed the recording of dendritic calcium
signals during visual stimulation (Denk and Detwiler, 1999; Euler
et al., 2002).
Uncaging and Photobleaching
Multi-photon photochemistry has been used to map receptor
sensitivities in single cells (Denk, 1994) and inside neural tissue
(Matsuzaki et al., 2001; Kasai et al., 2002).
Autofluorescence
Because MPM easily reaches into UV transition energies, it has
increasingly been used to study biological autofluorescence such
as from NADH (Piston et al., 1995; Piston and Knobel, 1999),
serotonin in living cells (Maiti et al., 1997), skin (Masters et al.,
1997), muscle cells (Schilders and Gu, 1999), glutathione in arabidopsis (Meyer and Fricker, 2000), mast cell secretion using 3P
excitation of serotonin (Williams et al., 1999), arctic fungus
(Arcangeli et al., 2000), collagen (Agarwal et al., 2001), biofilm
(Neu et al., 2002), tryptophan in proteins (Lippitz et al., 2002),
and flavoproteins (Huang et al., 2002). Recently, the sources of
autofluorescence from living tissue have been analyzed in more
detail (Zipfel et al., 2003) (see also Chapter 21, this volume).
545
Developmental Biology
Because of the superior depth penetration and the localized excitation associated with MPM, this approach has proven useful in
many developmental biological applications. Lineage tracing has
been performed using two-photon photorelease of caged fluorophores in sea urchin embryos (Summers et al., 1996; Piston
et al., 1998). Cellular and subcellular dynamics have been imaged
and measured using MPM during development of sea urchin
embryos (Summers et al., 1993, 1996), cell fusion in C. elegans
(Mohler et al., 1998; Periasamy et al., 1999), mammalian embryos
(Squirrell et al., 1999), zebrafish (Huang et al., 2001), and birds
(Dickinson et al., 2002).
In Vivo (Intact Animal) Imaging
In intact animals, the need for tissue penetration is maximal. High
resolution optical imaging inside living whole animals has therefore become the almost exclusive domain of two-photon
microscopy not only for functional calcium imaging (Svoboda et
al., 1997, 1999), but also to image blood flow in the fine capillaries (Kleinfeld et al., 1998; Chaigneau et al., 2003), gene expression and angiogenesis (Brown et al., 2001), and even the dynamics
of Alzheimer’s disease pathologies (Christie et al., 1998, 1999,
2001; Backskai et al., 2001) and, having previously been applied
to observe changes in dendrite structure in brain slices (Engert
and Bonhoeffer, 1999; Maletic-Savatic et al., 1999), two-photon
microscopy has most recently been used to study the long-term
dynamics of neuronal fine structure (Grutzendler et al., 2002;
Trachtenberg et al., 2002) in living animals.
OUTLOOK
Multi-photon excitation microscopy has extended the range of
laser scanning fluorescence microscopy especially where dynamic
imaging in living specimens is needed. Much progress has been
made in solving many of the technical impediments that existed in
the early days of MPM. Still, only a few of the many potential contrast mechanisms established for nonlinear optical spectroscopy
have been used for imaging purposes. This is mainly due to the
fact that often only a small number of photons can be collected
from each volume element in the small amount of time that the
beam dwells on each location. Increasing use is being made,
however, of second harmonic generation (Moreaux et al., 2001)
and Raman scattering (Zumbusch et al., 1999; Potma et al., 2002;
Volkmer et al., 2002) (see also Chapter 33, this volume.)
Phototoxicity in cells is still not well understood in general and
for ultrashort pulse illumination in particular. But the main limitation to even more widespread use of multi-photon excitation is not
due to fundamental physical, chemical, or biological problems, but
to the price and complexity of the instrumentation.
ACKNOWLEDGMENTS
The authors’ research underlying this chapter was sponsored by
the Developmental Resource for Biophysical Imaging and Optoelectronics at Cornell University (NIH-9 P41 EB001976 and
NSF-DIR-8800278), the Material Science Center Computing
Facility (NSF-DMR-9121564), other grants from NIH (DK53434,
CA86283) and NSF (BIR-98-71063), Lucent Technologies, Bell
Labs, and the Max-Planck Society.
546
Chapter 28 • W. Denk et al.
REFERENCES
Adronov, A., Frechet, J.M.J., He, G.S., Kim, K.S., Chung, S.J., Swiatkiewicz,
J., and Prasad, P.N., 2000, Novel two-photon absorbing dendritic structures, Chem. Mater. 12:2838–2841.
Agarwal, A., Coleno, M.L., Wallace, V.P., Wu, W.Y., Sun, C.H., Tromberg, B.J.,
and George, S.C., 2001, Two-photon laser scanning microscopy of epithelial cell-modulated collagen density in engineered human lung tissue,
Tissue Eng. 7:191–202.
Akaaboune, M., Grady, R.M., Turney, S., Sanes, J.R., and Lichtman, J.W.,
2002, Neurotransmitter receptor dynamics studied in vivo by reversible
photo-unbinding of fluorescent ligands, Neuron 34:865–876.
Albota, M., Beljonne, D., Bredas, J.L., Ehrlich, J.E., Fu, J.Y., Heikal, A.A.,
Hess, S.E., Kogej, T., Levin, M.D., Marder, S.R., McCord-Maughon, D.,
Perry, J.W., Rockel, H., Rumi, M., Subramaniam, C., Webb, W.W., Wu,
X.L., and Xu, C., 1998a, Design of organic molecules with large twophoton absorption cross sections, Science 281:1653–1656.
Albota, M.A., Xu, C., and Webb, W.W., 1998b, Two-photon fluorescence
excitation cross sections of biomolecular probes from 690 to 960 nm,
Appl. Opt. 37:7352–7356. http://www.drbio.cornell.edu/Infrastructure/
FluorescentProbes_WWW/CommerciallyAvailable.htm;
http://www.
drbio.cornell.edu/Infrastructure/NonlinearMicroscopies_WWW/vit.htm;
http://www.drbio.cornell.edu/Infrastructure/NonlinearMicroscopies_
WWW/3P/3PE.htm).
Andresen, V., Egner, A., and Hell, S.W., 2001, Time-multiplexed multifocal
multiphoton microscope, Opt. Lett. 26:75–77.
Arcangeli, C., Yu, W., Cannistraro, S., and Gratton, E., 2000, Two-photon autofluorescence microscopy and spectroscopy of antarctic fungus: New
approach for studying effects of UV-B irradiation, Biopolymers 57:
218–225.
Art, J.J., and Goodman, M.B., 1993, Rapid-scanning confocal microscopy,
Methods Cell Biol. 38:47–77.
Ashkin, A., Dziedzic, J.M., and Yamane, T., 1987, Optical trapping and
manipulation of single cells using infrared laser beams, Nature 330:769–
771.
Backskai, B.J., Kajdasz, S.T., Christie, R.H., Carter, C., Games, D., Seubert,
P., Schenk, D., and Hyman, B.T., 2001, Imaging of amyloid-beta deposits
in brains of living mice permits direct observation of clearance of plaques
with immunotherapy, Nat. Med. 7:369–372.
Beaurepaire, E., and Mertz, J., 2002, Epifluorescence collection in two-photon
microscopy, Appl. Opt. 41:5376–5382.
Beaurepaire, E., Oheim, M., and Mertz, J., 2001, Ultra-deep two-photon fluorescence excitation in turbid media, Opt. Commun. 188:25–29.
Bindhu, C.V., Harilal, S.S., Kurian, A., Nampoori, V.P.N., and Vallabhan,
C.P.G., 1998, Two and three photon absorption in rhodamine 6G methanol
solutions using pulsed thermal lens technique, J. Nonlinear Opt. Phys.
Mater. 7:531–538.
Birge, R.R., 1979, A theoretical analysis of the two-photon properties of linear
polyenes and the visual chromophores, J. Chem. Phys. 70:165–177.
Birge, R.R., 1986, Two-photon spectroscopy of protein-bound chromophores,
Acc. Chem. Res. 19:138–146.
Birge, R.R., and Zhang, C.-F., 1990, Two-photon double resonance spectroscopy of bacteriorhodopsin. Assignment of the electronic and dipolar
properties of the low-lying 1A*g+-like and 1B*u+like p,p* states, J. Chem.
Phys. 92:7178–7195.
Booth, M.J., and Hell, S.W., 1998, Continuous wave excitation two-photon fluorescence microscopy exemplified with the 647-nm ArKr laser line,
J. Microsc. 190:298–304.
Brakenhoff, G.J., Squier, J., Norris, T., Bliton, A.C., Wade, M.H., and Athey,
B., 1996, Real-time two-photon confocal microscopy using a femtosecond, amplified Ti : sapphire system, J. Microsc. 181:253–259.
Brown, E.B., Campbell, R.B., Tsuzuki, Y., Xu, L., Carmeliet, P., Fukumura, D.,
and Jain, R.K., 2001, In vivo measurement of gene expression, angiogenesis and physiological function in tumors using multiphoton laser scanning microscopy, Nat. Med. 7:866–870.
Bruchez, M. Jr., Moronne, M., Gin, P., Weiss, S., and Alivisatos, A.P., 1998,
Semiconductor nanocrystals as fluorescent biological labels, Science
281:2013–2016.
Chaigneau, E., Oheim, M., Audinat, E., and Charpak, S., 2003, Two-photon
imaging of capillary blood flow in olfactory bulb glomeruli, Proc. Natl.
Acad. Sci. USA 100:13081–13086.
Cheung, E.C., and Liu, J.M., 1991, Efficient generation of ultrashort, wavelenght-tunable infrared pulses, J. Opt. Soc. Am. B 8:1491–1506.
Christie, R.H., Bacskai, B.J., Zipfel, W.R., Williams, R.M., Kajdasz, S.T.,
Webb, W.W., and Hyman, B.T., 2001, Growth arrest of individual senile
plaques in a model of Alzheimer’s disease observed by in vivo multiphoton
microscopy, J. Neurosci. 21:858–864.
Christie, R.H., Zipfel, W.R., Williams, R.M., Webb, W.W., and Hyman, B.T.,
1998, Multiphoton imaging of Alzheimer’s disease neuropathology,
J. Neuropathol. Exp. Neurol. 57:145.
Christie, R.H., Zipfel, W.R., Williams, R.M., Webb, W.W., and Hyman, B.T.,
1999, In vivo multiphoton imaging of amyloid deposition in transgenic
mice, J. Neuropathol. Exp. Neurol. 58:204.
Corrie, J.E.T., and Trentham, D.R., 1993, Caged nucleotides and neurotransmitters, Bioorgan. Photochem. 2:243–305.
Curley, P.F., Ferguson, A.I., White, J.G., and Amos, W.B., 1992, Application
of a femtosecond self-sustaining mode-locked Ti-sapphire laser to the field
of laser scanning confocal microscopy, Opt. Quantum Electron.
24:851–859.
Davidovits, P., and Egger, M.D., 1969, Scanning laser microscope, Nature
233:831.
Debarbieux, F., Audinat, E., and Charpak, S., 2003, Action potential propagation in dendrites of rat mitral cells in vivo, J. Neurosci. 23:5553–
5560.
Denk, W., 1994, Two-photon scanning photochemical microscopy — Mapping
ligand-gated ion-channel distributions, Proc. Natl. Acad. Sci. USA
91:6629–6633.
Denk, W., 1996, Two-photon excitation in functional biological imaging,
J. Biomed. Opt. 1:296–304.
Denk, W., 2001, Optical beam power controller using a tiltable birefringent
plate, US Patent no. 6249379.
Denk, W., and Detwiler, P.B., 1999, Optical recording of light-evoked calcium
signals in the functionally intact retina, Proc. Natl. Acad. Sci. USA
96:7035–7040.
Denk, W., and Svoboda, K., 1997, Photon upmanship: Why multiphoton
imaging is more than a gimmick, Neuron 18:351–357.
Denk, W., Aksay, E., Baker, R., and Tank, D.W., 1997, Long term imaging of
[Ca++] in zebrafish embryos using two-photon microscopy, Soc. Neurosci.
Abstr. 646.
Denk, W., Delaney, K.R., Gelperin, A., Kleinfeld, D., Strowbridge, B.W., Tank,
D.W., and Yuste, R., 1994, Anatomical and functional imaging of neurons
using two-photon laser scanning microscopy, J. Neurosci. Methods 54:
151–162.
Denk, W., Strickler, J.H., and Webb, W.W., 1990, Two-photon laser scanning
fluorescence microscopy, Science 248:73–76.
Denk, W., Sugimori, M., and Llinas, R., 1995, Two types of calcium response
limited to single spines in cerebellar Purkinje cells, Proc. Natl. Acad. Sci.
USA 92:8279–8282.
Denk, W., Yuste, R., Svoboda, K., and Tank, D.W., 1996, Imaging calcium
dynamics in dendritic spines, Curr. Opin. Neurobiol. 6:372–378.
Dickinson, M.E., Murray, B.A., Haynes, S.M., Waters, C.W., and Longmuir,
K.J., 2002, Using electroporation and lipid-mediated transfection of GFPexpressing plasmids to label embryonic avian cells for vital confocal and
two-photon microscopy, Differentiation 70:172–180.
Dyba, M., and Hell, S.W., 2002, Focal spots of size lambda/23 open up farfield fluorescence microscopy at 33 nm axial resolution, Phys. Rev. Lett.
88:163901.
Eggeling, C., Widengren, J., Rigler, R., and Seidel, C.A.M., 1998, Photobleaching of fluorescent dyes under conditions used for single-molecule
detection: Evidence of two-step photolysis, Anal. Chem. 70:2651–2659.
Egner, A., and Hell, S.W., 2000, Time multiplexing and parallelization in multifocal multiphoton microscopy, J. Opt. Soc. A 17:1192–1201.
Egner, A., Jakobs, S., and Hell, S.W., 2002, Fast 100-nm resolution threedimensional microscope reveals structural plasticity of mitochondria in
live yeast, Proc. Natl. Acad. Sci. USA 99:3370–3375.
Engert, F., and Bonhoeffer, T., 1999, Dendritic spine changes associated with
hippocampal long-term synaptic plasticity, Nature 399:66–70.
Multi-Photon Molecular Excitation in Laser-Scanning Microscopy • Chapter 28
Euler, T., Detwiler, P.B., and Denk, W., 2002, Directionally selective calcium
signals in dendrites of starburst amacrine cells, Nature 418:845–852.
Fan, G.Y., Fujisaki, H., Miyawaki, A., Tsay, R.K., Tsien, R.Y., and Ellisman,
M.H., 1999, Video-rate scanning two-photon excitation fluorescence
microscopy and ratio imaging with cameleons, Biophys. J. 76:2412–
2420.
Fittinghoff, D.N., Schaffer, C.B., Mazur, E., and Squier, J.A., 2001, Timedecorrelated multifocal micromachining and trapping, IEEE J. Select
Topics Quantum Electron. 7:559–566.
Fork, R.L., Martinez, O.E., and Gordon, J.P., 1984, Negative dispersion using
pairs of prisms, Opt. Lett. 9:150–152.
Friedrich, D.M., and McClain, W.M., 1980, Two-photon molecular spectroscopy, Ann. Rev. Phys. Chem. 31:559–577.
Fu, Q., Mak, G., and van Driel, H.M., 1992, High-power, 62fs infrared optical
parametric oscillator synchronously pumped by a 76-MHz Ti:sapphire
laser, Opt. Lett. 17:1006–1010.
Furuta, T., Wang, S.S.H., Dantzker, J.L., Dore, T.M., Bybee, W.J., Callaway,
E.M., Denk, W., and Tsien, R.Y., 1999, Brominated 7-hydroxycoumarin4-ylmethyls: Photolabile protecting groups with biologically useful crosssections for two photon photolysis, Proc. Natl. Acad. Sci. USA 96:
1193–1200.
Gauderon, R., and Sheppard, C.J.R., 1999, Effect of a finite-size pinhole on
noise performance in single-, two-, and three-photon confocal fluorescence
microscopy, Appl. Opt. 38:3562–3565.
Goeppert-Mayer, M., 1931, Ueber Elementarakte mit zwei Quantenspruengen,
Ann. Phys. 9:273.
Goldstein, S.R., Hubin, T., and Smith, T.G., 1992, An improved no-movingparts video-rate confocal microscope, Microsc. Microanal. 23:437–446.
Gosnell, T.R., and Taylor, A.J., eds., 1991, Selected Papers on Ultrafast Laser
Technology, SPIE, Bellingham.
Grutzendler, J., Kasthuri, N., and Gan, W.B., 2002, Long-term dendritic spine
stability in the adult cortex, Nature 420:812–816.
Guild, J.B., Xu, C., and Webb, W.W., 1997, Measurement of group delay dispersion of high numerical aperture objective lenses using two-photon
excited fluorescence, Appl. Opt. 36:397–401.
Han, M.Y., Gao, X.H., Su, J.Z., and Nie, S., 2001, Quantum-dot-tagged
microbeads for multiplexed optical coding of biomolecules, Nat. Biotechnol. 19:631–635.
Hanninen, P.E., Lehtela, L., and Hell, S.W., 1996, Two- and multiphoton excitation of conjugate-dyes using a continuous wave laser, Opt. Commun.
130:29–33.
Hanninen, P.E., Soini, E., and Hell, S.W., 1994, Continuous-wave excitation
two-photon fluorescence microscopy, J. Microsc. 176:222–225.
He, G.S., Xu, G.C., Prasad, P.N., Reinhardt, B.A., Bhatt, J.C., and Dillard,
A.G., 1995, Two-photon absorption and optical-limiting properties of
novel organic-compounds, Opt. Lett. 20:435–437.
He, G.S., Yuan, L.X., Cheng, N., Bhawalkar, J.D., Prasad, P.N., Brott, L.L.,
Clarson, S.J., and Reinhardt, B.A., 1997, Nonlinear optical properties of
a new chromophore, J. Opt. Soc. Am. B 14:1079–1087.
Hell, S., and Stelzer, E.K.H., 1992, Fundamental improvement of resolution
with a 4pi-confocal fluorescence microscope using two-photon excitation,
Opt. Commun. 93:277–282.
Hell, S., Reiner, G., Cremer, C., and Stelzer, E.K.H., 1993, Aberrations in confocal fluorescence microscopy induced by mismatches in refractive index,
J. Microsc. 169:391–405.
Hell, S.W., and Andresen, V., 2001, Space-multiplexed multifocal nonlinear
microscopy, J. Microsc. 202:457–463.
Hell, S.W., Booth, M., Wilms, S., Schnetter, C.M., Kirsch, A.K., Arndt-Jovin,
D.J., and Jovin, T.M., 1998, Two-photon near- and far-field fluorescence
microscopy with continuous-wave excitation, Opt. Lett. 23:1238–1240.
Hellwarth, R., and Christiansen, P., 1974, Nonlinear optical microscopy examination of structure in polycrystalline ZnSe, Opt. Commun. 12:318–
322.
Helmchen, F., Fee, M.S., Tank, D.W., and Denk, W., 2001, A miniature headmounted two-photon microscope. High-resolution brain imaging in freely
moving animals, Neuron 31:903–912.
Helmchen, F., Tank, D.W., and Denk, W., 2002, Enhanced two-photon excitation through optical fiber by single-mode propagation in a large core, Appl.
Opt. 41:2930–2934.
547
Hockberger, P.E., Skimina, T.A., Centonze, V.E., Lavin, C., Chu, S., Dadras,
S., Reddy, J.K., and White, J.G., 1999, Activation of flavin-containing oxidases underlies light-induced production of H2O2 in mammalian cells,
Proc. Natl. Acad. Sci. USA 96:6255–6260.
Hopt, A., and Neher, E., 2001, Highly nonlinear photodamage in two-photon
fluorescence microscopy, Biophys. J. 80:2029–2036.
Huang, H., Vogel, S.S., Liu, N., Melton, D.A., and Lin, S., 2001, Analysis of
pancreatic development in living transgenic zebrafish embryos, Mol. Cell
Endocrinol. 177:117–124.
Huang, S.H., Heikal, A.A., and Webb, W.W., 2002, Two-photon fluorescence
spectroscopy and microscopy of NAD(P)H and flavoprotein, Biophys.
J. 82:2811–2825.
Kaiser, W., and Garrett, C.B.G., 1961, Two-photon excitation in CaF2 : Eu2+,
Phys. Rev. Lett. 7:229–231.
Kao, J.P.K., and Adams, S.R., 1993, Photosensitive caged compounds, In:
Optical Microscopy, Emerging Methods and Applications (B. Herman
and J.J. Lemasters, eds.), Academic Press, San Diego, California, pp. 27–85.
Kasai, H., Matsuzaki, M., and Ellis-Davies, G.C.R., 2002, Two-photon
mapping of functional glutamate receptors in dendritic spines of hippocampal CA1 pyramidal neurons, Jpn. J. Pharmacol. 88:S22.
Keller, U., 1994, Ultrafast all-solid-state laser technology, Appl. Phys. B Lasers
Opt. 58:347–363.
Keller, U., 1996, Materials and new approaches for ultrashort pulse lasers, Curr.
Opin. Solid State Mater. Sci. 1:218–224.
Kennedy, S.M., and Lytle, F.E., 1986, p-Bis(i-methylstyryl)benzene as a powersquared sensor for two-photon absorption measurements between 537 and
694 nm, Anal. Chem. 58:2643–2647.
Kim, O.K., Lee, K.S., Woo, H.Y., Kim, K.S., He, G.S., Swiatkiewicz, J., and
Prasad, P.N., 2000, New class of two-photon-absorbing chromophores
based on dithienothiophene, Chem. Mater. 12:284.
Kiskin, N.I., Chillingworth, R., McCray, J.A., Piston, D., and Ogden, D., 2002,
The efficiency of two-photon photolysis of a “caged” fluorophore, o-1-(2nitrophenyl)ethylpyranine, in relation to photodamage of synaptic terminals, Eur. Biophys. J. Biophys. Lett. 30:588–604.
Kleinfeld, D., Mitra, P.P., Helmchen, F., and Denk, W., 1998, Fluctuations and
stimulus-induced changes in blood flow observed in individual capillaries
in layers 2 through 4 of rat neocortex, Proc. Natl. Acad. Sci. USA
95:15741–15746.
Kliger, D.S.E., 1983, Ultrasensitive Laser Spectroscopy, Academic Press, New
York.
Kloppenburg, P., Zipfel, W.R., Webb, W.W., and Harris-Warrick, R.M., 2000,
Highly localized Ca2+ accumulation revealed by multiphoton microscopy
in an identified motoneuron and its modulation by dopamine, J. Neurosci.
20:2523–2533.
Koester, H.J., and Sakmann, B., 1998, Calcium dynamics in single spines
during coincident pre- and postsynaptic activity depend on relative timing
of back-propagating action potentials and subthreshold excitatory postsynaptic potentials, Proc. Natl. Acad. Sci. USA 95:9596–9601.
Koester, H.J., Baur, D., Uhl, R., and Hell, S.W., 1999, Ca2+ fluorescence
imaging with pico- and femtosecond two-photon excitation: Signal and
photodamage, Biophys. J. 77:2226–2236.
Konig, K., Becker, T.W., Fischer, P., Riemann, I., and Halbhuber, K.J., 1999,
Pulse-length dependence of cellular response to intense near-infrared laser
pulses in multiphoton microscopes, Opt. Lett. 24:113–115.
Konig, K., Kimel, S., and Berns, M.W., 1996, Photodynamic effects on human
and chicken erythrocytes studied with microirradiation and confocal laser
scanning microscopy, Lasers Surg. Med. 19:284–298.
Lansford, R., Bearman, G., and Fraser, S.E., 2001, Resolution of multiple green
fluorescent protein color variants and dyes using two-photon microscopy
and imaging spectroscopy, J. Biomed. Opt. 6:311–318.
Larson, D.R., Zipfel, W.R., Williams, R.M., Clark, S.W., Bruchez, M.P., Wise,
F.W., and Webb, W.W., 2003, Water-soluble quantum dots for multiphoton
fluorescence imaging in vivo, Science 300:1434–1436.
Lechleiter, J.D., Lin, D.T., and Sieneart, I., 2002, Multi-photon laser scanning
microscopy using an acoustic optical deflector, Biophys. J. 83:2292–
2299.
Lippitz, M., Erker, W., Decker, H., van Holde, K.E., and Basche, T., 2002, Twophoton excitation microscopy of tryptophan-containing proteins, Proc.
Natl. Acad. Sci. USA 99:2772–2777.
548
Chapter 28 • W. Denk et al.
Liu, T.M., Chu, S.W., Sun, C.K., Lin, B.L., Cheng, P.C., and Johnson, I., 2001,
Multiphoton confocal microscopy using a femtosecond Cr:forsterite laser,
Scanning 23:249–254.
Loudon, R., 1983, The Quantum Theory of Light, Oxford University Press,
New York.
Lozovoy, V.V., Pastirk, I., Walowicz, K.A., and Dantus, M., 2003, Multiphoton
intrapulse interference. II. Control of two- and three-photon laser induced
fluorescence with shaped pulses, J. Chem. Phys. 118:3187–3196.
Maiman, T.H., 1960, Stimulated optical radiation in ruby, Nature 187:493–494.
Mainen, Z.F., Maletic-Savatic, M., Shi, S.H., Hayashi, Y., Malinow, R., and
Svoboda, K., 1999a, Two-photon imaging in living brain slices, Methods
18:231–239.
Mainen, Z.F., Malinow, R., and Svoboda, K., 1999b, Synaptic calcium transients in single spines indicate that NMDA receptors are not saturated,
Nature 399:151–155.
Maiti, S., Shear, J.B., Williams, R.M., Zipfel, W.R., and Webb, W.W., 1997,
Measuring serotonin distribution in live cells with three-photon excitation,
Science 275:530–532.
Maletic-Savatic, M., Malinow, R., and Svoboda, K., 1999, Rapid dendritic morphogenesis in CA1 hippocampal dendrites induced by synaptic activity,
Science 283:1923–1927.
Marder, S.R., Torruellas, W.E., BlanchardDesce, M., Ricci, V., Stegeman, G.I.,
Gilmour, S., Bredas, J.L., Li, J., Bublitz, G.U., and Boxer, S.G., 1997,
Large molecular third-order optical nonlinearities in polarized carotenoids,
Science 276:1233–1236.
Masters, B.R., ed., 2003, Selected Papers on Multi-Photon Excitation
Microscopy, SPIE, Bellingham.
Masters, B.R., So, P.T.C., and Gratton, E., 1997, Multiphoton excitation fluorescence microscopy and spectroscopy of in vivo human skin, Biophys.
J. 72:2405–2412.
Matsuzaki, M., Ellis-Davies, G.C.R., Nemoto, T., Miyashita, Y., Iino, M., and
Kasai, H., 2001, Dendritic spine geometry is critical for AMPA receptor
expression in hippocampal CA1 pyramidal neurons, Nat. Neurosci. 4:
1086–1092.
McClain, W.M., 1971, Excited state symmetry assignment through polarized
two-photon absorption studies of fluids, J. Chem. Phys. 55:2789–2796.
Meyer, A.J., and Fricker, M.D., 2000, Direct measurement of glutathione in
epidermal cells of intact Arabidopsis roots by two-photon laser scanning
microscopy, J. Microsc. 198:174–181.
Milburn, T., Matsubara, N., Billington, A.P., Udgaonkar, J.B., Walker, J.W.,
Carpenter, B.K., Webb, W.W., Marque, J., Denk, W., McCray, J.A., and
Hess, G.P., 1989, Synthesis, photochemistry, and biological-activity of
a caged photolabile acetylcholine-receptor ligand, Biochemistry 28:49–
55.
Mohler, W.A., Simske, J.S., Williams-Masson, E.M., Hardin, J.D., and White,
J.G., 1998, Dynamics and ultrastructure of developmental cell fusions in
the Caenorhabditis elegans hypodermis, Curr. Biol. 8:1087–1090.
Moreaux, L., Sandre, O., Charpak, S., Blanchard-Desce, M., and Mertz, J.,
2001, Coherent scattering in multi-harmonic light microscopy, Biophys.
J. 80:1568–1574.
Mortensen, O.S., and Svendsen, E.N., 1981, Initial and final molecular states
as “virtual” states in two-photon processes, J. Chem. Phys. 74:3185–3189.
Muller, M., Squier, J., Wolleschensky, R., Simon, U., and Brakenhoff, G.J.,
1998, Dispersion pre-compensation of 15 femtosecond optical pulses for
high-numerical-aperture objectives, J. Microsc. 191:141–150.
Nakamura, O., 1993, Three-dimensional imaging characteristics of laser scan
fluorescence microscopy — two-photon excitation vs single-photon excitation, Optik 93:39–42.
Neu, T.R., Kuhlicke, U., and Lawrence, J.R., 2002, Assessment of fluorochromes for two-photon laser scanning microscopy of biofilms, Appl.
Environ. Microbiol. 68:901–909.
Neuman, K.C., Chadd, E., Liou, G.F., Brau, A., Bergman, K., and Block, S.M.,
1999, Characterization of photodamage induced by optical traps, Biophys.
J. 76:A96.
Nielsen, T., Frick, M., Hellweg., D., and Andresen, P., 2001, High efficiency
beam splitter for multifocal multiphoton microscopy, J. Microsc.
201:368–376.
Niggli, E., Piston, D.W., Kirby, M.S., Cheng, H., Sandison, D.R., Webb, W.W.,
and Lederer, W.J., 1994, A confocal laser scanning microscope designed
for indicators with ultraviolet excitation wavelengths, Am. J. Physiol.
266:C303–C310.
Oehring, H., Riemann, I., Fischer, P., Halbhuber, K.J., and Konig, K., 2000,
Ultrastructure and reproduction behaviour of single CHO-K1 cells
exposed to near-infrared femtosecond laser pulses, Scanning 22:263–
270.
Oertner, T.G., Sabatini, B.L., Nimchinsky, E.A., and Svoboda, K., 2002, Facilitation at single synapses probed with optical quantal analysis, Nat.
Neurosci. 10:10.
Ouzounov, D.G., Moll, K.D., Foster, M.A., Zipfel, W.R., Webb, W.W., and
Gaeta, A.L., 2002, Delivery of nanojoule femtosecond pulses through
large-core microstructured fibers, Opt. Lett. 27:1513–1515.
Patel, C.K.N., and Tam, A.C., 1981, Pulsed optoacoustic spectroscopy of condensed matter, Rev. Mod. Phys. 53:517–550.
Patterson, G.H., and Piston, D.W., 2000, Photobleaching in two-photon excitation microscopy, Biophys. J. 78:2159–2162.
Periasamy, A., Skoglund, P., Noakes, C., and Keller, R., 1999, An evaluation
of two-photon excitation versus confocal and digital deconvolution fluorescence microscopy imaging in Xenopus morphogenesis, Microsc. Res.
Technol. 47:172–181.
Piston, D.W., and Knobel, S.M., 1999, Real-time analysis of glucose metabolism by microscopy, Trends Endocrinol. Metab. 10:413–417.
Piston, D.W., Kirby, M.S., Cheng, H.P., Lederer, W.J., and Webb, W.W., 1994,
Two-photon-excitation fluorescence imaging of 3-dimensional calciumion activity, Appl. Opt. 33:662–669.
Piston, D.W., Masters, B.R., and Webb, W.W., 1995, 3-dimensionally resolved
Nad(P)H cellular metabolic redox imaging of the in-situ cornea with twophoton excitation laser-scanning microscopy, J. Microsc. 178:20–27.
Piston, D.W., Sandison, D.R., and Webb, W.W., 1992, Time-resolved fluorescence imaging and background rejection by two-photon excitation in laser
scanning microscopy, Proc. SPIE. 1640:379–389.
Piston, D.W., Summers, R.G., and Webb, W.W., 1993, Observation of nuclear
division in living sea urchin embryos by two-photon fluorescence
microscopy, Biophys. J. 63:A110.
Piston, D.W., Summers, R.G., Knobel, S.M., and Morrill, J.B., 1998, Characterization of involution during sea urchin gastrulation using two-photon excited
photorelease and confocal microscopy, Microsc. Microanal. 4:404–414.
Potma, E.O., Jones, D.J., Cheng, J.X., Xie, X.S., and Ye, J., 2002, Highsensitivity coherent anti-Stokes Raman scattering microscopy with two
tightly synchronized picosecond laser, Opt. Lett. 27:1168–1170.
Powers, P.E., Tang, C.L., and Cheng, L.K., 1994, High-repetition-rate femtosecond optical parametric oscillator based on Cstioaso4, Opt. Lett. 19:
37–39.
Rehms, A.A., and Callis, P.R., 1993, Two-photon fluorescence excitationspectra of aromatic-amino-acids, Chem. Phys. Lett. 208:276–282.
Sabatini, B.L., and Svoboda, K., 2000, Analysis of calcium channels in single
spines using optical fluctuation analysis, Nature 408:589–593.
Sandison, D.R., and Webb, W.W., 1994, Background rejection and signal-tonoise optimization in confocal and alternative fluorescence microscopes,
Appl. Opt. 33:603–615.
Schilders, S.P., and Gu, M., 1999, Three-dimensional autofluorescence spectroscopy of rat skeletal muscle tissue under two-photon excitation, Appl.
Opt. 38:720–723.
Schönle, A., and Hell, S.W., 1998, Heating by absorption in the focus of an
objective lens, Opt. Lett. 23:325–327.
Sheppard, C.J.R., and Cogswell, C.J., 1990, Confocal microscopy with detector arrays, J. Mod. Opt. 37:267–279.
Sheppard, C.J.R., and Cogswell, C.J., 1991, Effects of aberrating layers and
tube length on confocal imaging properties, Optik 87:34–38.
Sheppard, C.J.R., and Gu, M., 1990, Image-formation in two-photon fluorescence microscopy, Optik 86:104–106.
Sheppard, C.J.R., and Gu, M., 1994, Imaging performance of confocal fluorescence microscopes with finite-sized source, J. Mod. Opt. 41:1521–1530.
Sheppard, C.J.R., and Kompfner, R., 1978, Resonant scanning optical microscope, Appl. Opt. 17:2879–2882.
Singh, S., and Bradley, L.T., 1964, Three-photon absorption in napthalene crystals by laser excitation, Phys. Rev. Lett. 12:612.
Spence, D.E., Kean, P.N., and Sibbett, W., 1991, 60-fsec pulse generation from
a self-mode-locked Ti:sapphire laser, Opt. Lett. 16:42–44.
Multi-Photon Molecular Excitation in Laser-Scanning Microscopy • Chapter 28
Squier, J.A., Fittinghoff, D.N., Barty, C.P.J., Wilson, K.R, Muller, M., and
Brakenhoff, C.J., 1998, Characterization of femtosecond pulses focused
with high numerical aperture optics using interferometric surface-thirdharmonic generation, Opt. Commun. 147:153–156.
Squirrell, J.M., Wokosin, D.L., White, J.G., and Bavister, B.D., 1999, Longterm two-photon fluorescence imaging of mammalian embryos without
compromising viability, Nat. Biotechnol. 17:763–767.
Stelzer, E.H.K., Hell, S., and Lindek, S., 1994, Nonlinear absorption extends
confocal fluorescence microscopy into the ultra-violet regime and confines
the illumination volume, Opt. Commun. 104:223–228.
Straub, M., and Hell, S.W., 1998, Multifocal multiphoton microscopy: A fast
and efficient tool for 3-D fluorescence imaging, Bioimaging 6:177–185.
Summers, R.G., Morrill, J.B., Leith, A., Marko, M., Piston, D.W., and
Stonebraker, A.T., 1993, A stereometric analysis of karyokinesis, cytokinesis and cell arrangements during and following 4th cleavage period in
the sea-urchin, Lytechinus variegatus, Dev. Growth Diff. 35:41–57.
Summers, R.G., Piston, D.W., Harris, K.M., and Morrill, J.B., 1996, The orientation of first cleavage in the sea urchin embryo, Lytechinus variegatus,
does not specify the axes of bilateral symmetry, Dev. Biol. 175:177–183.
Svoboda, K., and Block, S.M., 1994, Biological applications of optical forces,
Ann. Rev. Biophys. Biomol. Struct. 23:247–285.
Svoboda, K., Denk, W., Kleinfeld, D., and Tank, D.W., 1997, In vivo dendritic
calcium dynamics in neocortical pyramidal neurons, Nature 385:161–165.
Svoboda, K., Denk, W., Knox, W., and Tsuda, S., 1996a, Two-photon laser
scanning fluorescence microscopy of living neurons using a diode-pumped
Cr : LiSrAlFl laser mode-locked with a saturable Bragg reflector, Opt. Lett.
21:1411–1413.
Svoboda, K., Helmchen, F., Denk, W., and Tank, D.W., 1999, Spread of dendritic excitation in layer 2/3 pyramidal neurons in rat barrel cortex in vivo,
Nat. Neurosci. 2:65–73.
Svoboda, K., Tank, D.W., and Denk, W., 1996b, Direct measurement of coupling between dendritic spines and shafts, Science 272:716–719.
Tan, Y.P., Llano, I., Hopt, A., Wurriehausen, F., and Neher, E., 1999, Fast scanning and efficient photodetection in a simple two-photon microscope,
J. Neurosci. Methods 92:123–135.
Theer, P., Hasan, M.T., and Denk, W., 2003, Two-photon imaging to a depth
of 1000 mm in living brains by use of a Ti : Al2O3 regenerative amplifier,
Opt. Lett. 28:1022–1024.
Tirlapur, U.K., Konig, K., Peuckert, C., Krieg, R., and Halbhuber, K.J., 2001,
Femtosecond near-infrared laser pulses elicit generation of reactive
oxygen species in mammalian cells leading to apoptosis-like death, Exp.
Cell Res. 263:88–97.
Trachtenberg, J.T., Chen, B.E., Knott, G.W., Feng, G., Sanes, J.R., Welker, E.,
and Svoboda, K., 2002, Long-term in vivo imaging of experiencedependent synaptic plasticity in adult cortex, Nature 420:788–794.
Valdmanis, J.A., and Fork, R.L., 1986, Design considerations for a femtosecond pulse laser balancing self phase modulation, group-velocity dispersion, saturable absorption, and saturable gain, IEEE J. Quant. Electron.
22:112–118.
Ventelon, L., Moreaux, L., Mertz, J., and Blanchard-Desce, M., 1999, New
quadrupolar fluorophores with high two-photon excited fluorescence,
Chem. Commun. 20:2055–2056.
Ventelon, L., Moreaux, L., Mertz, J., and Blanchard-Desce, M., 2002, Optimization of quadrupolar chromophores for molecular two-photon absorption, Synth. Met. 127:17–21.
Visser, T.D., Brakenhoff, G.J., and Groen, F.C.A., 1991, The one-point fluorescence response in confocal microscopy, Optik 87:39–40.
549
Volkmer, A., Book, L.D., and Xie, X.S., 2002, Time-resolved coherent antiStokes Raman scattering microscopy: Imaging based on Raman free
induction decay, Appl. Phys. Lett. 80:1505–1507.
Walowicz, K.A., Pastirk, I., Lozovoy, V.V., and Dantus, M., 2002, Multiphoton
intrapulse interference. 1. Control of multiphoton processes in condensed
phases, J. Phys. Chem. A 106:9369–9373.
Wang, S.S.H., Denk, W., and Hausser, M., 2000, Coincidence detection in
single dendritic spines mediated by calcium release, Nat. Neurosci.
3:1266–1273.
Whinnery, J.R., 1974, Laser measurement of optical absorption in liquids, Acc.
Chem. Res. 7:225–231.
Williams, R.M., Shear, J.B., Zipfel, W.R., Maiti, S., and Webb, WW., 1999,
Mucosal mast cell secretion processes imaged using three-photon
microscopy of 5-hydroxytryptamine autofluorescence, Biophys. J.
76:1835–1846.
Williams, S.A., and Callis, P.R., 1990, Two-photon electronic spectra of
nucleotides, Proc. SPIE 1204:332–343.
Wilson, T., and Sheppard, C., 1984, Theory and Practice of Scanning Optical
Microscopy, Academic Press, New York.
Wokosin, D.L., Centonze, V., White, J.G., Armstrong, D., Robertson, G.,
and Ferguson, A.I., 1996a, All-solid-state ultrafast lasers facilitate multiphoton excitation fluorescence imaging, IEEE J. Quant. Electron.
2:1051–1065.
Wokosin, D.L., Centonze, V.E., Crittenden, S., and White, J., 1996b, Threephoton excitation fluorescence imaging of biological specimens using an
all-solid-state laser, Bioimaging 4:1–7.
Wolleschensky, R., Feurer, T., Sauerbrey, R., and Simon, I., 1998, Characterization and optimization of a laser-scanning microscope in the femtosecond regime, Appl. Phys. B Lasers Opt. 67:87–94.
Xu, C., and Denk, W., 1997, Two-photon optical beam induced current imaging
through the backside of integrated circuits, Appl. Phys. Lett. 71:
2578–2580.
Xu, C., and Denk, W., 1999, Comparison of one- and two-photon optical beaminduced current imaging, J. Appl. Phys. 86:2226–2231.
Xu, C., and Webb, W.W., 1996, Measurement of two-photon excitation cross
sections of molecular fluorophores with data from 690 to 1050 nm, J. Opt.
Soc. Am. B 13:481–491.
Xu, C., Guild, J., Webb, W.W., and Denk, W., 1995, Determination of absolute
two-photon excitation cross-sections by in-situ 2nd-order autocorrelation,
Opt. Lett. 20:2372–2374.
Xu, C., Zipfel, W., Shear, J., Williams, R., and Webb, W., 1996, Multiphoton
fluorescence excitation: New spectral windows for biological nonlinear
microscopy, Proc. Natl. Acad. Sci. USA 93:10763–10768.
Yuste, R., and Denk, W., 1995, Dendritic spines as basic functional units of
neuronal integration, Nature 375:682–684.
Zhang, Q., Piston, D.W., and Goodman, R.H., 2002, Regulation of corepressor function by nuclear NADH, Science 295:1895–1897.
Zipfel, W.R., Williams, R.M., Christie, R., Nikitin, A.Y., Hyman, B.T., and
Webb, W.W., 2003, Live tissue intrinsic emission microscopy using multiphoton-excited native fluorescence and second harmonic generation,
Proc. Natl. Acad. Sci. USA 100:7075–7080.
Zojer, E., Beljonne, D., Kogej, T., Vogel, H., Marder, S.R., Perry, J.W., and
Bredas, J.L., 2002, Tuning the two-photon absorption response of
quadrupolar organic molecules, J. Chem. Phys. 116:3646–3658.
Zumbusch, A., Holtom, G.R., and Xie, X.S., 1999, Three-dimensional vibrational imaging by coherent anti-Stokes Raman scattering, Phys. Rev. Lett.
82:4142–4145.