Univerza v LjubljaniFakulteta za matematiko in fiziko

Oddelek za fiziko

Seminar II - 2. letnik, II. stopnja

Two-photon fluorescence microscopy

Author: Boštjan Kokot

Advisor: Janez Štrancar

Coadvisor: Iztok Urbancic

Ljubljana, March 2015

Abstract

Two-photon laser scanning microscopy is one of the well established tools for studying biological sys-tems, enabling in vivo and in situ imaging. Here the basics for understanding of fluorescence microscopyare presented, starting with single-photon fluorescence. It is followed by a brief explanation of confocalmicroscopy. Next the extension to two-photon excitation is explained and two-photon microscopy ispresented. Further, a comparison between confocal and two-photon microscopy is given. Finally, theextension to three-photon microscopy is presented, alongside with a few neat examples of multi-photonimaging.

Contents

1 Introduction 1

2 Single-photon fluorescence 2

3 Confocal microscopy 3

4 Photobleaching and photodamage 4

5 Two-photon microscopy 5

6 Comparison between two-photon and confocal single-photon microscopy 7

7 Multi-photon microscopy 8

8 Use of multi-photon microscopy 8

9 Conclusion 10

References 10

1 Introduction

With the development of optical microscopy in vivo and in situ studies of microscopic biological samples,such as living cells, have been made possible [1]. First biological samples were observed under thetransmission optical microscope. Thus acquired images suffer from a serious lack of contrast due to smalldifferences in light abosrption within watery cells (Fig. 1a). A much better contrast can be achieved byadding fluorescent markers to the sample and filtering the scattered incident light out of the image, thussubtracting the background (Fig. 1b).

Fluorescent markers can be embedded into the structures we wish to study. This enables us a muchdeeper insight into the detailed structure of the sample than the conventional transmission microscopy(Fig. 1). With the invention of the confocal microscopy, the much anticipated three-dimensional live celland tissue imaging has been enabled but was unfortunately limited to optically ’thin’ samples [2]. Asan alternative the two-photon microscopy was developed, offering great depth resolution and penetrationdepth of 1-2 mm due to longer excitation wavelengths being used [3]. With infrared excitation theemission is confined to 1 µm cube, thus better axial resolution is achieved. Furthermore the phototoxicityand photodamage are reduced, which leads to the much searched for extended viability (survival) of livesamples [4]. The development of stable solid-state lasers with tunable wavelengths has enabled accessto the two-photon microscopy to a wider community of researchers [3].

Two-photon fluorescence has been extended to multi-photon fluorescence with even better depthresolution and excitation confinement [5]. It has provided us with a tool for three-dimensional imaging.

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(a) (b)

Figure 1: (a) An example of transmission optical microscope image of human cancer cell lacking in con-trast. One can barley discern the borders of three cells. (b) An example of fluorescence microscopy imageof the same sample as in panel (a). The contrast of the outer membranes is enhanced with fluorescentmarker.

2 Single-photon fluorescence

If one is to understand the principle of two-photon excitation he has to first understand the basic principleof single-photon excitation. It can easily be explained with a Jablonsky diagram represented in Fig. 2a.

When a fluorophore, a molecule that is able to fluoresce, is illuminated with the light of the wave-length in the range of absorption spectra (Fig. 3b, blue line), it absorbs a photon. Photon absorptionexcites an electron from the ground state to the first electronic state [6, 7, 8]. It relaxes to the lowestexcited state through molecular vibrational states before dropping back to the ground state. During theelectron transition to the ground state the molecule emits light of longer wavelength (lower energy). Theshift of the emission spectrum with respect to the absorption spectrum is presented in Fig. 2b known asthe Stokes shift. This enables us to filter the scattered excitation light and thus only measure the emittedfluorescence. Consequently, we only obtain high-contrast images where the bright labeled structuresappear on a dark background. Most of the fluorophores used in fluorescence microscopy absorb light inultraviolet (UV), visible or infrared part of the spectrum.

The probability for single-photon fluorescence is linear in the flux of incident Φ. We can thus writethe rate W1 at which the fluorescence occurs as

W1 = σΦ, (1)

where σ is the single-photon absorption cross section, which has the units of surface [1]. The crosssection is a measure for the probability for the absorption [9].

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(a) (b)

Figure 2: (a) The jablonski diagram, showing the intramolecular processes of fluorescence. Absorbinga photon, an electron is excited the to first electronic state. After a relaxation through the molecularvibrational states it drops back to the ground state, emitting a photon of a longer wavelength [7]. (b) Anexample of absorption and emission spectra for a typical fluorophore with the Stoke’s shift labeled inbetween [10].

3 Confocal microscopy

While with the single-photon focusing of the excitation light deeper into the sample is made possible,we cannot avoid exciting out-of-focus planes in the process. Consequently the signal from the lower andupper planes will blur our image [3]. We can filter the excitation light to attain a much better contrastcompared to transmission microscopy, but this technique does not offer us a fine depth sectioning of thesample.

An elegant solution is the confocal microscopy, schematics for the setup shown in Fig. 3a. We limitthe excitation light to one focal plane with a pinhole [11, 12]. Another, or the same, pinhole is used tolimit the light emitted from the same plane in the sample. With this method we gather the signal fromone point of the plane. To acquire a full two-dimensional image we have to either raster-scan the wholeplane by deflecting a laser beam (upper inlay Fig. 3a) or use a confocal Nipkow disk with a wide-fieldsource (lower inlay of Fig. 3a). Raster-scanning is achieved by two synchronized galvanometric mirrors.The setup is similar to the one presented in Fig. 3a with the source being laser and mirrors positionedright after the first pinhole. On the other hand, the holes of the confocal disc represent a set of pinholesthrough which more points of the plane are illuminated and recorded simultaneously. We acquire thewhole image by rotating the disc.

In order to obtain as much depth resolution as possible, the pinhole size must be diffraction lim-ited [11, 13]. Smaller pinhole leads to better contrast but less gathered signal. Laser-scanning versionsoften offer an adjustable pinhole, whereas a fixed compromise is implemented in Nipkow disk setups.

Confocal microscopy improves the depth sectioning of our sample compared to single-photon flu-orescence microscopy, but it has one crucial weakness for imaging of photosensitive fluorophores. Aswe also excite fluorophores in out-of-focus planes, with excitation limited to a cone-shaped pattern, rep-resented in Fig. 3b (left). That means that before we acquire the image of that plane photosensitivefluorophores will degrade.

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Lastly, we gather much less signal while using the pinhole because we are able to detect only lightscattered directly towards the detector. Emission light from focal planes deeper in the sample scattersmore, therefore less light is scattered directly towards the detector, meaning more light is rejected bythe pinhole. This consequently leads to much smaller penetration depth compared to single-photonmicroscopy.

(a) (b)

Figure 3: (a) Schematics of an experimental setup for confocal microscopy. The pinhole after the sourcerestricts the excitation to the focal plane (green beam) and the one in front of the detector rejects theout of focus light (red beam). The inlays show two versions of scanning the sample plane also knownas raster-scan and a confocal (Nipkow) disc used for wide-source microscopy [11]. (b) Comparisonbetween single-photon confocal and two-photon excitation [14].

4 Photobleaching and photodamage

Imaging of biological samples labeled with fluorophores is unavoidably accompanied by photobleachingand photodamage. While the first is concerned with fluorophores and their basic properties the other isconnected with live-sample properties.

Molecule with the electron in the excited state is more reactive. Interaction with the chemical envi-ronment causes the cleaving of covalent bonds [15]. Consequently molecule loses the ability to fluoresce,what is known as photobleacing. More sensitive flurophores bleach after only few absorption-emissioncycles while the robust ones can survive trillions. Photobleaching occurs gradually and is observed assignal fading through time of sample exposure (Fig. 4).

Although photobleaching is unwanted with microscopy techniques and can be partially countered bymeasuring the time-lapse images and correcting images accordingly, it can present us with informationabout local environment of the fluorophore[6]. Problem occurs when one would like to study complexdynamical structures that are composed of multiple components. The systems with more components areunpredictable, so their evolution after loss of signal, cannot be described using mathematical algorithms.

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Figure 4: An example of photobleaching in imaging of a typical mammalian cell. Flourescence signal(top row) was measured at the beginning of the experiment, after 3 s and 30 s with corresponding imagesof regular microscopy (bottom row). Loss of the signal after only 30 s is evident. The third bright fieldimage has alteration of the membrane due to photodamage pointed out with the arrow [16].

In the real experimental conditions (sparse labeling, fluorophores with low cross-sections, short ex-posure times for imaging fast processes, etc.) images with insufficient signal-to-noise ratio are obtained.As Eq. 1 suggests, the signal can be improved by increasing the photon flux Φ.

While increasing the intensity when dealing with non-biological samples might be fine, biologicalsamples are very sensitive to energy dissipation and temperature changes. If we heat the living sample toomuch, processes that keep it alive will be terminated. Light can also trigger oxidative chemical reactions,byproducts of which consequently alter their homeostasis (natural balance), potentially leading to theirdeath [16, 17]. In conclusion, keeping the photodamage at the minimum extends the viability of cells,which is crucial for the acquisition of biologically relevant data.

5 Two-photon microscopy

We can extend the idea of single-photon excitation by simultaneously exciting the fluorophore withtwo photons instead of one. This was theoretically first suggested by Maria Göppert-Mayer in early1930’s [18] and experimentally confirmed in 1960’s by Kaiser et. al [19]. It has remained unexploiteduntil 1990, when Denk et al. showed the possibility of exciting a 1 µm thick slice on the depth of∼ 100 µm, using two-photon excitation [2].

In two-photon microscopy we use two photons of approximately twice the wavelength of single-photon excitation to excite the fluorophore to the same excited state as before [7, 20]. Emission processhappens in the same way as with the one-photon excitation. Schematics similar to the one for single-photon excitation in Fig. 2a is shown in Fig. 5a.

There are two direct consequences to the fact that we need two photons to excite a molecule. Thefirst is a higher intensity of light needed for fluorescence to happen as the fluorophore is able to absorbtwo photons only if they hit the molecule at approximately the same time. The probability for that tooccur is much lower than for the single-photon excitation, thus the need for a high-power light source.

The second consequence is the square flux dependency of fluorescence rate W2 that can be written as

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W2 = δΦ2, (2)

where δ represents the cross section for two-photon absorption measured in units of Göppert-Mayer,1 GM = 10−50 cm4s = 10−58 m4s [1, 4, 17]. Due to square flux dependence we excite only the fluo-rophores in the focus as only there enough photons are crowded in a small enough place. This can beachieved with the combination of high numerical aperture optics and high laser peak powers of 1 W.Sources able to achieve high peak powers are picosecond pulsed lasers, operating at the 100 MHz fre-quency. High numerical aperture enables us to achieve submicron slice thickness, which allows us betterlight focusing and consequently the lower powers of lasers. Results of simulation of beam spreading de-pendent on laser power is presented in Fig. 5b. We can see, that with higher laser powers beam spreadsmore, thus lowering the axial resolution of slice thickness. Comparison to single-photon excitation whereout of focus planes are excited is presented in Fig. 3b.

(a) (b)

Figure 5: (a) The Jablonski diagram for two-photon excitation. It represents the same process as inFig. 2a, the only difference being the absorption of two photons of double the wavelength [7]. (b) Asimulation of a Gaussian laser beam profile for different numerical apertures (NA) of objective lenses.Contours represent different intensities. To each simulation, the spatial scale in µm is added. We canobserve the axial spreading of the beam and lower in-plane spatial resolution for lower aperture lenses [4].

The strange units of δ can be explained in terms of one-photon cross section. We say that a moleculeis in a virtual state after absorbing one lower energy photon waiting to absorb another one. The energiesof a pair of photons can only combine if those photons happen to pass through a small area ∆A in ashort time interval ∆t. That means that another photon has to arrive to the molecule, while the excitedelectron still persist in a virtual state. It follows that δ ≈ ∆A ∆t σ, σ being the cross section for single-photon absorbtion. We can make a simplified estimation of the cross section, neglecting parity rules andintermediate molecular states. The spatial scale in which electronic interactions take places is ∆A ≈ σ ≈10−16 cm2 and the time scale is determined by a typical lifetime of a molecular virtual state ∆t ≈ 10−17 s.If we combine the values above we get a two-photon cross section δ ≈ 10−49 cm4s = 10 GM [1]. Typicalvalues of the cross-section for the useful fluorophores are from 10−1 GM to 102 GM, meaning that ourcoarse estimation agrees with the experiment. Sometimes we are forced to use a fluorophore with lower

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cross sections and therefore need to use higher laser powers. That affects the axial resolution as well asshortens the viabilty of our sample.

To achieve the two-photon excitation high-power source is needed. While the continuous-wave lasersare possible, they have high average power, which results in the high damage to the biological samples.Therefore the continuous-wave lasers are rather avoided. Instead we use the pulsed lasers with high peakpowers (1 W) and low average powers (50 mW), realized by short pulses (≈ 100 ps) at repetition rates≈ 80 Mhz [1, 2, 4, 7]. The repetition rate must not exceed the inverse fluorescence lifetime (typically1-10 ns or 108-109 Hz) as the fluorophores become saturated. The saturation causes the loss of signalbecause the fluorophores cannot relax to the ground state before we try to excite them again [2, 7].

Signal can be optimized with tuning the excitation wavelength to match the absorption spectra. Thecloser to the absorption peak we excite a fluorophore, the higher the emitted fluorescence and strongerthe signal will be. This can be achieved with wavelength tunable lasers. The present choice of source isTi:Sapphire solid-state laser tunable in the range 700 nm - 1000 nm [1, 2]. The other potential sourceis Nd:YAG laser, tunable in range 600 nm - 800 nm [1]. Besides some solid state lasers for the longerwavelength range are being developed.

We can further optimize the signal with the high numerical aperture to achieve the highest localizationof the excitation possible. This also results in better depth resolution as can be seen from the simulationof the effect of different numerical aperture lenses on the laser beam profile represented in Fig. 5b. Wecan see that the lower the aperture the more elongated the shape of the beam in axial direction will be [4].To achieve the desired axial micron resolution we have to therefore use the high numerical aperture lenses(NA ≈ 1.2).

Lastly we can optimize the signal by removing the pinhole, which can be done because the fluores-cence excitation is very localized. Removing the pinhole affects the spatial resolution, but enables usto penetrate much deeper into the sample. Imaging with the pinhole is called the descanned detectionand it proceeds in the same way as the already described confocal microscopy imaging [21]. Scanningwithout the pinhole also known as the non-descanned detection. It is done in the same way as confocalmicroscopy with the pinhole removed and the external detector with a larger field-of-view added to thesetup. We need to place the external detector beside the sample to gather not only the light scattereddirectly backwards, but to gather us much light as possible scattered to the other directions as well. Thedetected light is then assigned to the recored focal point, from which we know it has emitted. This ispossible only because of the very localized nature of the two-photon excitation [21].

6 Comparison between two-photon and confocal single-photon microscopy

There are many merits to using the two-photon microscopy over the confocal and one significant down-side. The first advantage of two-photon over confocal microscopy is its micron depth resolution, allowingfine three-dimensional sectioning, some fine examples presented later in this paper. To illustrate the abil-ity of the two-photon microscopy to section micron thick slices a micron thick bleaching pattern, whereone focal plane has been scanned repeatedly until sample has bleached, is presented in Fig. 6a [2].

Moreover, due to the longer wavelengths of the excitation light that observes only the excited areaof the sample, penetration depth into the sample is much greater than that of a confocal microscopy.The reason for such a difference is lower absorption and scattering of near-infrared light, with the higherwavelengths. Even more, at wavelengths between 700 nm and 1100 nm one-photon absorption is thelowest for major cellular absorbers (melanin, haemoglobin, water), meaning they are most transparentfor this range of wavelengths. That results in even lower photodamage and greater penetration depth thanit usually would [17]. Presented in Fig. 6c and Fig. 6d (from left to right) is depth illumination of samplefor confocal, descanned two-photon and nondescanned two-photon microscopy, respectively [22].

Another quality of the two-photon microscopy is much less bleaching compared to the confocal

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microscopy. Due to the very localized nonlinear in flux nature of the two-photon excitation the bleachingwill only occur in the plane of sample we wish to observe (Fig. 6a) and not in the cone shape as is thecase with the confocal microscopy (Fig. 6a) [2].

(a)

(b)

(c)

(d)

Figure 6: (a) Bleaching pattern after illuminating the sample in confocal single-photon microscopy fora long period of time. (b) Bleaching pattern after illuminating the sample for a long period of time intwo-photon microscopy. Contrast with confocal microscopy is obvious [2]. (c) Single photon confocalmicroscopy image showing penetration depth. Scale bar represents 20 µm [22]. (d) Two-photon des-canned and non-descanned image of the same sample as Fig. 6c, with the same scale bar. Comparisonbetween the three clearly shows superb resolution of two-photon microscopy.

As mentioned before, one shortcoming of two-photon microscopy is its lowered lateral spatial resolu-tion resulting from using the light with approximately twice the wavelength. Due to diffraction limit [23]single-photon confocal microscopy would reslove objects as twice small as two-photon microscopy [1],enough to study the details of cells and tissues. For even more demanding experiments, however, theso-called super resolution microscopes, also based on multi-photon excitation, have achieved resolutionunder the diffraction limit.

7 Multi-photon microscopy

The obvious extension of two-photon excitation leads to three-photon excitation. This requires the lightof approximately three times the wavelength and even higher intensities [8].

Advantages from using two photons instead of one scale the same way when we use three photonsinstead of one: better penetration depth, better localization of excitation and less photodamage, but lowerin-plane spatial resolution and even higher required peak intensities of laser [5, 17].

8 Use of multi-photon microscopy

Here I present some of the latest imaging achievements using multi-photon microscopy, but deeper un-derstanding can be found elsewhere [5, 17, 24, 25].

Multi-photon microscopy has been developed so far that we can perform imaging in vivo [5, 17]. Afew examples of resourceful use of two and three-photon microscopy are listed below. In Fig. 7a andFig. 7a a three-photon excitation three-dimensional images of veins in mouse brain and kidney bud are

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presented, respectively [26, 27]. Fig. 8a represents two-photon-excited autofluorescence of human skin atdifferent tissue depths with high spatial resolution [17]. Lastly, Fig. 8b holds a from in vivo three-photonmicroscopy three-dimensional reconstruction of mouse brain neurons through the depth of 1200 µm [5].

(a) (b)

Figure 7: (a) Three-dimensional model three-photon imaging reconstructed from the data of mouse brainveins. Field of view: 250 × 250 × 500 µm volume [26]. (b) Three-dimensional model two-photon imagereconstructed from the data of the kidney bud. Two buds are approximately 5 µm apart [27].

(a) (b)

Figure 8: (a) Two-photon microscopy images of autofluorescence of human skin, freshly obtained bybiopsy from breast cancer at different depths with great spatial resolution [17]. (b) A three dimensionalreconstruction of mouse brain neurons from three-photon microscopy performed in vivo [5].

The localization of multi-photon excitation and low average powers, compared to continuous wavelasers, have enabled perform nanosurgery on cells without excessive damage [17]. It also leads to awhole new series of experiments in micropharmacology [1]. UV light emitted after the multi-photonexcitation photoactivates the caged compound resulting in its photolysis. For example we can mimic therelease of neurotransmitter that occurs at the synapse [17].

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9 Conclusion

The basic concepts of fluorescent microscopy have been briefly explained, with two-photon microscopyexplained more in detail. Comparison has shown that two-photon laser scanning microscopy excels overconfocal microscopy. It enables deeper penetration into tissues and provides higher depth resolution.Localization of excitation and emission has eliminated the need for a pinhole, allowing to gather moresignal from the sample. Photodamage and phototoxicity are also confined to the focal point. The methodhas become accesible to a wider range of researchers with the development of new solid-state turn-keylasers [3]. Lastly, the expansion from two to three-photon microscopy has been presented, alongsidewith three excellent examples of multi-photon imaging.

References

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[11] I. Urbancic. Odziv biomembranskih domen na zunanje dražljaje. PhD, Faculty of Natural Sciencesand Mathematics, University of Maribor, 2013.

[12] James B. Pawley. Handbook of Biological Confocal Microscopy. Springer, New York, 2006.

[13] http://en.wikipedia.org/wiki/Diffraction-limited_system#The_Abbe_diffraction_limit_for_a_microscope. (3. 3. 2015).

[14] http://microscopy.berkeley.edu/courses/tlm/2P/index.html. (25. 2. 2015).

[15] https://en.wikipedia.org/wiki/Photobleaching. (23. 2. 2015).

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[20] M. Oheim, D. J. Michael, M. Geisbauer, D. Madsen, and R. H. Chow. Adv. Drug. Deliver. Rev.,58:789, 2006.

[21] http://www.microscopyu.com/articles/fluorescence/multiphoton/multiphotonintro.html. (3. 3. 2015).

[22] G. Brown and H. Hempstead. Technical notes. Bio-Rad, 5.

[23] E. Hecht. Optics (4th Edition), year =.

[24] A. Ustione and D. W. Piston. 243:221, 2011.

[25] K. Zagorovsky and W. C. W. Chan. Nat. Mater., 12:285, 2013.

[26] http://www.conoptics.com/neurosciencebiophotonics-brain-initiative-opportunities-optics-photonics/.(25. 2. 2015).

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