TPE in conjunction with laser scanning microscopy offers several advantages over conventional single-photon confocal laser scanning microscopy. One major advantage is its capability to provide optical sectioning with subcellular resolution from deeper within scattering biological specimens than single-photon confocal microscopy, making this technique particularly attractive for intravital fluorescence imaging. Systematic observations on a variety of biological samples have provided experimental evidence that TPE improves imaging penetration depth by at least 2-fold relative to confocal imaging.20–22 Several reasons account for the increase in penetration depth. First, the dependence of fluorophore excitation on the second power of laser light intensity confines photon absorption to a narrow region at the plane of focus, where photon flux is highest (see Equation 1 and Figure 1b). Thus, unlike single-photon confocal microscopy, TPE microscopy lacks linear absorption of the excitation beam by fluorophores above the plane of focus, which can significantly reduce excitation light before it reaches fluorophores within deeper tissue regions.20 Second, the longer wavelengths used for TPE are scattered by the tissue much less than the shorter wavelengths used for confocal microscopy,23 resulting in deeper penetration of the focused laser beam. Third, scattered light emitted from an excited fluorophore within the focal volume does not contribute to the final image in confocal microscopy because it is indistinguishable from fluorescent light generated in out-of-focus areas and is rejected by the pinhole in the emission path. By contrast, because TPE never generates out-of-focus fluorescence (see Figure 1b), scattered photons from fluorophore emission can be used to generate the TPE image, resulting in increased fluorescence collection efficiency and thus greater signal intensity at any given tissue depth.23 With common mode-locked Ti:Sapphire lasers, imaging depths in a variety of tissues are typically Oheim et al have extended the penetration depth of two-photon microscopy by 100 µm by means of a low-magnification, high numerical aperture objective lens.23 Theer et al were able to increase the reach of multiphoton microscopy to 1 mm in brain tissue by using regenerative amplification of 200-kHz pulses to achieve high peak powers while maintaining reasonable average powers.19 However, optical resolution degraded substantially at greater depths, possibly resulting from excitation saturation.12 Finally, Webb and colleagues recently showed that multiphoton microscopy through gradient index lenses enables subcellular resolution imaging several millimeters deep in the brain of anesthetized, intact animals.24 Importantly, the spatial resolution of TPE microscopy is maintained with increasing tissue depth,20,25 provided that excitation saturation is negligible.12 Although increasing incident average laser power is one common means to achieve greater penetration depth of the excitation beam,19,23,26 it also increases the probability of both excitation saturation and higher-order (more than second-order) excitation processes, resulting in decreased optical resolution and more extensive photodamage, respectively.14,15,17
Future strategies to improve tissue penetration depth include development and use of dyes having a larger two-photon absorption cross section, use of longer wavelengths of the excitation light, and measures to increase collection efficiency of the microscope (such as better transmittance of filters, beam splitters, etc). Because the fraction of scattered photons in the emitted fluorescence signal increases with increasing imaging depth, optics with a larger effective angular acceptance should markedly augment the contribution of scattered emission photons to the fluorescence image, thereby further improving depth resolution.23
In comparison with confocal imaging, TPE fluorescence microscopy reduces overall photobleaching and photodamage by limiting it to the narrow region around the focal plane.1 This reduction becomes important when collecting three-dimensional data sets (z-series) in thick specimens. By contrast, in confocal microscopy, all focal planes are exposed to excitation light every time an optical plane is collected. Therefore, confocal laser scanning fluorescence microscopy can cause more extensive photobleaching and photodamage than TPE fluorescence microscopy if z-stacks are taken repeatedly over time (time-lapse imaging). For endogenous (eg, nicotine adenine dinucleotide [NADH]) and exogenous fluorophores (eg, ion-sensitive indicators) that require UV excitation, conventional confocal imaging has only limited potential because of extensive UV-induced photodamage and photobleaching.27 Through the use of TPE microscopy, it has been possible to perform spatially resolved quantitative measurements of NADH levels and ion concentrations in single cells and intact tissue.28,29
Finally, the spatial confinement of TPE provides unprecedented opportunities for three-dimensionally localized photochemistry at a subfemtoliter scale, such as photoactivated release of caged calcium ions7–9 or neurotransmitters,30,31 as well as for measurements of three-dimensional mobility of fluorescent molecules using fluorescence photobleaching recovery.10,11
Optical resolution is linearly dependent on wavelength of the excitation light. It is therefore surprising that the difference in effective resolving power between near-infrared light TPE microscopy and visible light confocal microscopy has been shown to be much less than one would predict from the differences in excitation wavelength.13 This finding results from the nonlinear nature of the two-photon absorption process, which in turn limits the spread of the excitation volume, thereby preventing a significant decrease in resolution resulting from the use of longer-wavelength excitation light.