Multiphoton Microscopy

A. Neumann, Y. Kuznetsova

Introduction Multi-Photon Fluorescence Microscopy is a relatively novel imaging technique in cell biology. It relies on the quasi-simultaneous absorption of two or more photons (of either the same or different energy) by a molecule. With normal fluorescence microscopy, one cannot resolve deep structures within a specimen because of large depth of focus. An enhanced method, the confocal microscopy [1], allows scanning a series of thin optical `slices' through the thickness (Z-direction) of the specimen, and its three-dimensional representation can be generated and manipulated with image-processing software. But, because we are interested in observing living specimens, often at several stages during development, we run into some serious problems with normal confocal fluorescence microscopy. One of these is photobleaching of the fluorescent label (chromophore)[2]. Because the small confocal aperture blocks most of the light emitted by the tissue, including light coming from the plane of focus, the exciting laser must be very bright to allow an adequate signal-to-noise ratio. This bright light causes fluorescent dyes to fade within minutes of continuous scanning. In addition to photobleaching, phototoxicity is also a problem [3]. Excited fluorescent dye molecules generate toxic free-radicals. Thus, one must limit the scanning time or light intensity if one hopes to keep the specimen alive. The solution was: Two-Photon Microscopy.

Multi-photon absorption Multi-photon absorption was predicted by Maria Göppert-Mayer in 1930 , and the proof-of-principle was performed in the 1960s using continuous-wave laser sources [2]. Now multi-photon fluorescence microscopy in cell biology uses pulsed lasers (pulse width in the range of 20 fs to 5 ps, typically ~ 100 fs) with high repetition rates (10 - 100 MHz). Two-photon confocal imaging differs from one-photon confocal imaging in a number of ways (Fig 1).

Fig. 1. One and two-photon excitation.

In normal fluorescence microscopy, the fluorescent molecule absorbs a photon of light of a particular wavelength (usually around 350-500 nm) [4]. This photon of light 'carries' a particular amount of energy, the shorter the wavelength, the higher the associated energy. This energy then causes an electron in the molecule to be excited from the ground state into a higher energy level. Electrons quickly decay back to the ground state by losing the energy that they absorbed, in the form of new photons of light (fluorescence). In theory, this fluorescent light should be of the same wavelength as the photons that the molecule originally absorbed. In practice, however, the electron loses a bit of energy by decaying through the vibrational ladders of the higher energy level, resulting in fluorescent light of a longer wavelength than the absorbed light. In two-photon excitation, a fluorophore can be excited from its ground state by the near-simultaneous absorption of two photons of light of half the energy of its conventional one-photon absorption band. Fluorophores such as fura-2 and indo-1, which absorb in the UV to green region (350-500 nm), can therefore be excited by an ultra-fast, mode-locked Ti:sapphire laser, emitting light in the red to near infrared region of the spectrum (700-900 nm).

Absorption for one, two and three photon process is described by the following formulas correspondently:

dI dI dI = −αI , = −βI 2 , = −γI 3 dz dz dz where I – intensity, α,β,γ - absorption coefficients. As far as intensity of the focused beam is reciprocal proportional to the squared beam waist: 2   z   πnw2 w2 = w2 1+    , where z = 0 , λ - is wave length, z-distance from focal plane, 0   z   0 λ   0   then the absorption for one, two and three photon process has the following spatial dependence:

dI 1 dI 1 dI 1 ~ − , ~ − , ~ − dz 2 2 dz (z 2 + z 2 ) 2 dz (z 2 + z 2 )3 z0 + z 0 0

Two-Photon Microscopy We see that the probability of two-photon absorption increases quadratically with illumination intensity, and to the forth power with proximity to focal plane, what essentially limiting excitation to a tiny focal volume (Fig. 2a). Also this limits both out-of-focus background excitation and provides innate confocality, since only fluorophore in the plane of the focused beam will be excited. The main difference in confocal and two-photon detection is shown on Fig. 2b. Here the confocal one-photon excitation imaging compared with two-photon imaging in scattering tissue. Due to the longer wavelength, less excitation light is lost to scattering when using two-photon excitation. Ballistic (non-scattering) and diffusing fluorescence photons can be used in the two-photon case, but only ballistic photons can be used in the confocal case [5]. So, in multi-photon excitation more fluorescence photons are detected from a focal point than from a confocal method.

a) b)

Fig. 2 a) Scheme of two- and one-photon excitation, b) comparison between confocal and two-photon detection

Three-photon microscopy

Three-photon excitation can also be used in certain circumstances [6]. In this case three photons are absorbed simultaneously, effectively tripling the excitation energy. Using this technique, UV excited fluorophores may be imaged with IR excitation. Because excitation levels are dependent on the cube of the excitation power, resolution is improved (for the same excitation wavelength) compared to two-photon excitation where there is quadratic power dependence. It is possible to select fluorophores such that multiple labeled samples by can be imaged by combination of 2- and 3 photon excitation, using a single IR excitation source (Fig. 3).

Fig. 3 Combination 2-photon (red and green) and 3-photon (blue) image of C.elegans embryo

Scanningless depth resolved microscopy Multiphoton microscopes offer an own mechanism for optical sectioning. Rather than rejecting out-of-focus scattering, it is practically eliminated by its nonlinear dependence on the illumination intensity. A multiphoton process, is efficient only when the peak intensity of the illuminating light is high, i.e. at the focal spot [7]. Acquisition of an image still requires, however, scanning of either the sample or the laser beam, resulting in a similar restriction of the image frame rate. Numerous methods have been developed to increase the image acquisition rate in both confocal and multiphoton microscopes. Most involve multi-point illumination and scanning in a single spatial axis. Common examples are single-axis scanning and the use of line illumination [8, 9], or rotation of a patterned disk (in confocal microscopy) [10] or a lenslet array (in multiphoton microscopy) [11, 12, 13]. The same idea was implemented using beamsplitter arrays [14, 15]. Others have used chromatic multiplexing [16]. Overall, however, image acquisition times have not been reduced significantly beyond video-rate imaging. Performing depth-resolved microscopy completely without scanning could bring image acquisition times down to the shortest limit determined by the signal level. In a simple scanningless implementation of a multiphoton microscope the sample is illuminated with a low NA [14], but this results in a complete loss of the depth resolution property. It has been suggested that wide-field illumination maintaining the depth resolution can be done by massively splitting the excitation pulse with beamsplitter arrays, so that the generated foci overlap each other spatially but are temporally delayed relative to each other [15, 17, 18]. Up to several tens of such time-separated foci have been generated using beamsplitter arrays, but massive splitting as required for scanningless microscopy has never been practically realized, and may result in difficulty of achieving uniform illumination for the many thousands of beamlets required to image a reasonably sized area. Then another method was suggested to use temporal focusing of the excitation pulse rather than spatial focusing [19]. This principle is demonstrated in Fig. 4. The frame a) depicts the standard multiphoton technique, where an ultrashort pulse is spatially focused, generating high peak intensity at the focus; the temporal profile of the pulse (i.e. the pulse duration) remains nearly unchanged as it propagates through the sample. The panel b) is an illustration of the new method. The illumination beam excites the full frame of interest in the specimen, an area greater by orders of magnitude than the diffraction limited spot of the objective lens. The depth resolution is achieved by controlling the temporal profile of the pulse, which is compressed as it propagates through the sample, reaching its peak value at the focal plane, and stretching again as it propagates beyond it. The depth-resolved multiphoton signal can now be collected from the entire illuminated frame using an imaging setup. Two major advantages this scheme are an inherent high illumination uniformity as it is directly related to the spatial uniformity of the excitation pulse, and scalability of the illuminated area. The basic principle enabling these temporal manipulations can be best understood using the setup shown schematically in Fig. 4(c,d). It consists of a thin scattering plate, sitting at the front focal plane of a telescope. A short pulse with duration τ illuminates the plate, and each point scatters the light into many directions; each point in the image plane is also illuminated for duration τ, since all light rays emerging from a single point in the object travel identical optical pathlengths and reach the image at the same time, a concept known as the Fermat principle. However, any other point P, at some distance either from the scattering plate or from the image plane is illuminated for a longer duration, dictated by the different trajectories taken by the rays reaching it. A simple geometrical optics estimation of the illumination duration at the point P shows that the time delay of the illumination due to the two paths shown in the figure is ∆t = z(cos−1(θ)−1)/c. The further is P from the scatterer, the longer is the illumination time. Similarly, points in front or in the back of the focal plane see extended illumination. Thus, it is simple to arrive at: n − n 2 − NA 2 z ∆t ≈ ⋅ n 2 − NA 2 c where n is the refractive index and NA the numerical aperture.

a) c)

b) d)

Fig. 4. (a) Standard multiphoton microscopy scheme: a high peak intensity at the focus of the objective lens is generated by spatial focusing of the beam, while the temporal profile of the pulse remains unchanged in passing through the sample. (b) Scanningless method. The beam is weakly focused, covering an area which is many orders of magnitude larger than diffraction limited spot. High peak intensity at the focus is achieved by modification of the temporal profile of the pulse as it propagates, reaching its peak at the focus of the objective lens. (c) Principle of the scanningless TPEF microscope. A short pulse impinges upon a scatterer. At a point P further away, the pulse duration is longer due to the difference in the length of trajectories taken by the rays reaching it from different locations on the scatterer. Only at the image plane of the telescope is the pulse duration restored to its initial value, in accordance with the Fermat principle. (d) The experimental setup: The input beam impinges upon a grating, aligned perpendicular to the optic axis of the microscope. The grating is imaged through a high magnification telescope, comprised of an achromatic lens and the microscope objective, on the sample. Fluorescence is epi-detected and imaged onto a CCD using a dichroic mirror.

Equating the pulse duration τ with ∆t should result in an approximate expression for the depth of the temporal focus. For high NA lenses (assuming NA/n≈0.9) this results in z≈ τc/2, corresponding to about 3 µm for 20fs pulses. In general, to optimize the nonlinear effect, short pulses are advantageous. Too short pulses (<20 f s), however, should be avoided since they suffer from significant material dispersion and since their bandwidth may exceed the TPEF absorption linewidth. Long pulses (> 100 fs) increase complexity since large magnifications and a highly dispersive grating are required. Blazed gratings are well suited for use as scattering plates, due to their low cost, availability, and extremely high efficiency of off-axis scattering. In order to measure a maximal signal, the peak intensity at the focal plane should be as high as possible, yet remain below the damage threshold of biological specimen [9]. So full-frame depth-resolved multiphoton microscopy is possible, using an extremely simple setup and standard components, but Z-resolution will decrease in compare to spatial focusing procedure.

Conclusions Advantages of multiphoton imaging:

• Confocal imaging suffers a deterioration in signal-to-background when obtaining images from deep within a sample. Multiphoton imaging is largely immune from these effects as little fluorescence is generated away from the point of illumination (Fig 5) [6]. The confocal image shows a significant increase in local background resulting in a lower contrast image. However, the multiphoton image maintains contrast even at significant depths within a light scattering sample.

Fig. 5. Comparison between confocal and multiphoton imaging

• Multi-photon fluorescence microscopy allows imaging in high absorbing media. • Increased detection sensitivity of fluorophores by reducing autofluorescence or background. • All the emitted photons from multi-photon excitation can be used for imaging (in principle) therefore no confocal blocking apertures or descanning optics have to be used allowing for a simpler, more light efficient optical design. • Unlike confocal or wide-field fluorescence imaging, multiphoton imaging only excites fluorophore in the plane of the optical section being imaged. Therefore there is no fluorophore excitation in the bulk of the specimen. This considerably reduces the generation of phototoxic products. • UV fluorophores may be excited using a lens that is not corrected for UV. • A scanningless setup can enable full-frame video-rate fluorescence lifetime imaging

Limitations of multi-photon excitation:

• Slightly lower resolution with a given fluorophore when compared to confocal imaging. This loss in resolution can be eliminated by the use of a confocal aperture at the expense of a loss in signal. • Thermal damage can occur in a specimen if it contains chromophores that absorb the excitation wavelengths, such as the pigment melanin. • Only works with fluorescence imaging. • Currently rather expensive • Slow imaging due to scanning process by spatial focusing / low Z-resolution by temporal focusing

References and links

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