Since its initial demonstration in 1990 [1
], multi-photon microscopy (MPM) has become one of the most powerful imaging modalities in diverse areas of biology, including neuroscience, immunology, embryology and cancer research [2
]. Because of its highly spatially confined ‘nonlinear’ excitation, MPM does not suffer from emission generated out of focal volume and can achieve sub-micrometer, three-dimensional (3D) optical resolution without a detector pinhole. Since virtually all of the emission light contributes to useful signal, the detection systems for MPM usually have high efficiency that enhances imaging at larger depths in optically turbid media. In addition, the long excitation wavelengths used in MPM allow deeper penetration into biological tissue and produce less damage to specimens [4
]. In addition to two-photon fluorescence excitation (2PE) [2
], which is the most commonly used in MPM, other nonlinear processes have been applied for imaging including: second harmonic generation (SHG) [10
], coherent anti-Stokes Raman scattering [17
], three-photon fluorescence excitation (3PE) [18
] and third harmonic generation (THG) [22
One of the main limiting factors preventing widespread use of MPM is the high cost and complexity of short pulse laser sources [4
]. The price of the laser system can constitute up to 50% of the total cost of commercial multi-photon microscopes, and the percentage could be significantly higher for the custom-built microscopes [4
]. In the following paragraphs we will briefly describe the requirements for the laser sources in MPM and the reasons for their high cost. More details on this subject can be found in [2
Several parameters of the laser, such as average power (Pave), pulsewidth (τ), pulse repetition rate (f), and central wavelength (λexc), are critical for the quality of the MPM image. Due to the nonlinear nature of excitation in MPM, the average emission signal intensity is proportional to Pnave/(τf)n−1, where n is the order of the nonlinear process. One can ideally increase the average power of the laser in order to increase the average emission signal intensity; however, the average power that can be applied to a biological sample is limited by destructive thermal effects, photo-toxicity and photobleaching.
Although it has been demonstrated that using very short pulses (~10 fs) may be beneficial for MPM [31
], pulse durations of 50–100 fs offer several practical advantages [2
]. It is challenging to compensate for the effects of group delay dispersion (GDD) of the microscope optics for very short pulses, since higher order dispersion effects become dominant for broad bandwidths. Without careful dispersion precompensation a 10-fs pulse may even produce a weaker fluorescent signal than a 100-fs pulse [31
]. Also, laser sources generating 50–100 fs are more common than 10-fs lasers. Finally, the spectrum of a very short pulse may be wider than the absorption spectrum of the fluorophore. This reduces the excitation efficiency and can also make it difficult to separate emission from excitation light.
Higher pulse repetition rates of the laser source may increase the fluorescence emission rate and reduce the required scan time, if the average power is not a limiting factor. However, most fluorophores used in MPM have an upper state lifetime of ≈5 ns, and for very high repetition rates (f > 200 MHz) absorption saturation may limit the obtainable signal strength [30
] and decrease the resolution due to accumulation of fluorophores in the excited state. A pulse repetition rate of ~100 MHz is close to an optimum choice for many MPM applications [8
]; for historical reasons, most lasers used in MPM have a repetition rate of 80 MHz.
The excitation wavelengths for MPM are typically in the near-infrared spectral region. In particular, two-photon absorption (2PA) cross sections of the most common fluorophores are centered between 700 nm and 1000 nm and are usually broader and blue-shifted with respect to corresponding one-photon spectra, with full-width half-maximum (FWHM) bandwidths typically exceeding 100 nm [8
]. Therefore, it is usually not critical to tune the excitation wavelength to the exact maximum of the 2PA cross section when imaging a single fluorophore. However, wavelength tunability may be very important for imaging using several fluorophores in the same sample [38
]. The maximum two-photon ‘action’ cross section for common fluorophores is in the range of 1–300 GM, and depending on this cross section, only a few tens of milliwatts or less of the excitation power at the focus of high numerical aperture (NA) objective is sufficient to saturate the excitation, assuming 100 MHz pulse repetition rate and 100 fs pulsewidth [1
]. Propagation of the optical beam through the microscope optics and focusing deep into the scattering biological tissue produce attenuation from absorption and scattering as well as optical aberrations, which reduce the excitation efficiency within the focal volume. Investigators typically compensate for this reduced efficiency by increasing the laser power. For a laser producing 50–100 fs pulses at a repetition rate of ~100 MHz, an average power of 200 mW is sufficient for most MPM applications [2
provides a representative list of laser sources that have been used for MPM. In its early years, MPM imaging was performed using colliding pulse dye lasers [1
]: the only source at the time that could generate sub-picosecond pulses with sufficient peak powers, at repetition rates of ~100 MHz [30
]. However, due to their complexity as well as to maintenance issues [2
], dye lasers are difficult to use for MPM. The real breakthrough in MPM came with the development of Ti:Sapphire lasers, which are much more powerful and user friendly than dye lasers [2
]. Today’s commercial Ti:Sapphire lasers can generate sub-100-fs pulses with hundreds of kW peak power and are broadly tunable (680 to 1080 nm) [43
]. The tunability can be extended further to the visible and mid-infrared regions using optical nonlinear schemes such as optical parametric oscillators (OPOs), photonic crystal fibers (PCFs) and second harmonic generation (SHG) [46
]. The majority of today’s MPM systems use fully automated Ti:Sapphire lasers, where the central wavelength and average power of the laser are adjusted under computer control [43
]. The main disadvantage of Ti:Sapphire femtosecond laser technology is its high cost. Ti:Sapphire gain medium requires pumping in the green spectral region. Unfortunately, direct diode pumping is not available at these wavelengths, and pump sources use the second harmonic generated by the diode-pumped Nd-doped laser systems. These Nd-based systems are bulky and a standard pump laser costs in the range of $50–100k, making the overall cost of the Ti:Sapphire lasers and the MPM systems quite high, thus limiting widespread adoption of MPM [2
Representative list of laser sources that have been used for multi-photon microscopy.
Beyond the popular Ti:Sapphire laser, several other lasers have been demonstrated for MPM. One alternative is Neodymium or Ytterbium based systems that can generate ~200-fs pulses with ~100 kW of peak power at ~1050 nm wavelengths [30
]. The main disadvantage of these rare earth element lasers is their narrow emission bandwidth, which prevents tunability. Most fluorophores used in MPM have a TPE spectrum centered below 1000 nm [8
] and these laser sources cannot reach this part of the spectrum directly. However, as with Ti:Sapphire, nonlinear effects can be used to generate tunable outputs from these sources, at the expense of increased complexity and cost [50
Diode lasers, which are compact, easy to use and low-cost, represent another possible light source for MPM. Gain-switched diode lasers can provide picosecond pulses and by using external amplifiers, peak powers above ~1 kW level can be obtained [53
]. Repetition rate is easily adjustable by simply changing the current driving frequency, which might be advantageous in some MPM experiments. However, low peak powers and issues related to amplified-spontaneous-emission and tunability currently limit their effective use in MPM.
In this paper, we describe the Cr:LiCAF laser as a promising light source for MPM. Cr:LiCAF belongs to the family of Cr3+
-doped colquiriite solid-sate laser gain media, where the most well-known members are: Cr:LiSAF [59
], Cr:LiSGaF [65
], and Cr:LiCAF [66
]. Similar to Ti:Sapphire, Cr:Colquirites have broad emission bandwidths around 800 nm, enabling the generation of pulses as short as ~10 fs [69
]. The possibility of developing a low cost and efficient femtosecond technology based on diode pumped Cr:Colquirites was suggested more than five years ago with the demonstration of battery powered femtosecond operated Cr:LiSAF laser [61
]. The main advantage of Cr:Colquirites over Ti:Sapphire is that they can be directly diode-pumped by inexpensive diodes in the red region of the spectrum; hence, the total material costs of an entire laser system could be below ~$10k [27
]. We believe that, with the recent advances, Cr:Colquiriites lasers have the potential to become efficient and low-cost excitation light sources for MPM and could enable wide-spread use of MPM technology, by reducing cost.
Cr:Colquiriite lasers have previously been demonstrated for MPM imaging () [42
]. Using an 500-mW master oscillator power amplifier (SDL, Inc.) with a diffraction limited beam as the pump source, Svoboda et al. developed a Cr:LiSAF laser producing 90-fs pulses with ~0.3 nJ pulse energy at ~860 nm and a repetition rate of 150 MHz for MPM [42
]. However, this pump source is very expensive and the available peak power from the laser was below 5 kW, limiting deep tissue imagining in scattering media. A multimode diode pumped Cr:LiSAF laser with similar specifications of 100–200 pulse duration, 0.375 nJ pulse energy at 80 MHz repetition rate was also used to demonstrate MPM imaging; however, peak powers were limited [57
The Cr:LiCAF laser used in this study employs inexpensive single mode diodes, costing only ~$150 each, as the pump source and uses double-chirped mirrors for dispersion compensation. A semiconductor saturable absorber mirror (SESAM) [70
], also referred to as a saturable Bragg reflector (SBR) [71
], was used for mode locking, which enables self-starting and robust laser operation. The laser produces ~70 fs long pulses (FWHM), with 1.8-nJ of energy, centered around 800 nm wavelengths, with a repetition rate of 100 MHz. The average output power is 180 mW, and the peak power is as high as 25 kW. This is approximately an order of magnitude higher peak power than earlier Cr:Colquiriite lasers that were used for MPM, and the laser design is also simpler. These improvements over the previous lasers are the result of advances in high-power visible diodes, mirror coatings, SESAMs/SBRs and the quality of Cr:Colquiriite crystals.
This paper is organized as follows: Section 2 describes the Cr:LiCAF laser setup used in the MPM studies. Section 3 describes methods for representative examples of MPM imaging with the Cr:LiCAF laser in neuroscience, immunology, endocrinology and cancer research. Section 4 presents results and Section 5 discussion and conclusions.