Fully parallel optical methods proposed for 3D activation in an all-optical configuration have been presented for optogenetics and make use of extended light-pattern formation to cover the entire neuron soma. Light-patterns can be either large Gaussian beams generated by underfilling the objective numerical aperture (NA; low-NA Gaussian beams), or beams created with more flexible light-patterning methods such as CGH, generalized phase contrast (GPC) or amplitude modulation. As already mentioned, the limit of extended light patterns is the deterioration of the axial confinement, an issue that can be solved by using temporal focusing. Common experimental configurations of TF make use of a diffraction grating in a plane conjugate to the focal plane of the microscope objective (image formation plane), separating the spectral frequencies of the laser femtosecond pulses (dispersion of different wavelengths in different angles) (Oron et al., 2005; Zhu et al., 2005). In other words, the image projected at the grating plane is the image formed at the sample plane, while 2P absorption of light projected in any other plane before or after the grating, as it is the case when you generate a 3D pattern distribution, will be strongly weakened by the pulse broadening. This has, until recently, limited TF-light shaping to 2D configurations. Methods have been proposed for axial displacement of the TF plane for Gaussian beams, by introducing group velocity dispersion (Durst et al., 2006, 2008; Dana and Shoham, 2012). However, they are only suitable for remotely displacing one plane and they are not compatible with light-patterning techniques (Leshem et al., 2014), as they can displace the TF plane but not the spatial focusing one.

Hernandez et al. (2016) solved the problem, by introducing the axial displacement mechanism after the grating for TF, which decoupled lateral light shaping from axial displacements. The system used a conventional 2D-CGH with TF for lateral light-patterning and a second SLM placed at a Fourier plane after the grating to introduce the desired axial shift from the original focal plane, via a lens-effect phase modulation. This configuration also enabled the generation of different excitation patterns at distinct axial planes, by addressing the two SLMs in multiple regions, tiled vertically to the direction of dispersion for TF. With this configuration, researchers demonstrated for the first time the generation of multi-plane temporally focused patterns, reaching a volume of 240 × 240 × 260 μm³ with the axial confinement varying from 5 μm Full Width Half Maximum (FWHM) at the center of the field of excitation (FOE) to 10 μm at the edges of it, tested with spots of 20 μm in diameter. The number of regions equal to the number of planes to be addressed. The system was used for selective 2P 3D photoconversion of the Kaede protein (Isobe et al., 2010) in the brain of zebrafish larvae (photoconversion at 800 nm with 0.1–4.0 mW/μm² depending the illumination duration) and for 2D optogenetic activation of ChR2 in zebrafish spinal cord neurons co-expressing GCaMP5G (excitation at 900 nm with 0.6 mW/μm²). Monitoring of Ca²⁺ traces in that case was performed with visible illumination and two-color HiLo imaging (Lim et al., 2008) (at ∼30.8 μW/mm²).

Despite the flexibility of this system, an inherent limitation occurs in the maximum number of axial planes that could be addressed because of the physical tiling of the SLMs, before the quality of the holographic spots get distorted (∼6 planes) (Hernandez et al., 2016). Although this can be sufficient for a number of biological applications (Yang et al., 2016), three new studies, have recently proposed ways to increase both the number of planes and the FOE. They all used the same principle: having a beam-shaping method for creating a 2D temporally focused pattern and using a SLM at a Fourier plane of the TF system for lateral and axial multiplexing of this pattern in several positions in 3Ds via 3D-CGH (Figure 2A). 3D-CGH is used to generate arrays of 3D diffraction-limited spots, using variations of the Gerchberg-Saxton algorithm (Gerchberg and Saxton, 1972; Di Leonardo et al., 2007) to calculate the phase profile to address to the SLM.

In two of these works the light-shaping part is a TF-Gaussian beam (Pégard et al., 2017; Sun et al., 2018) (Figure 2B). In (Pégard et al., 2017) 3D scanless holographic optogenetics with temporal focusing (3D-SHOT) was used to generate a large number of temporally focused spots, each of them fitting the size of a cell soma of pyramidal neurons (i.e., ∼10–15 μm FWHM lateral size). With this approach, researchers reported the possibility to project hundreds of excitation spots in a total volume of 350 × 350 × 280 μm³. (Sun et al., 2018) presented a very similar approach as the one presented by (Pégard et al., 2017), using a TF experimental configuration with two gratings instead of a single grating and a lens. The beam’s lateral size was 2.5 μm in diameter and the axial FWHM of the 2P light intensity was 7.5 μm. This study did not include any biological demonstration. The authors used the generated spots in direct laser writing inside a glass, which relies on non-linear light absorption at the focus. Fabrication was conducted either on the surface or inside standard microscope glass slides. As a novelty, they showed that they were able to place focal points at high lateral proximity (4 m) with minimal interference between them. That is thanks to the pulse front tilt effect, a property inherent to TF systems where the arrival time of an ultrashort pulse in a certain plane varies across the beam profile thus creating a tilt between the pulse front and the direction perpendicular to the beam. In this way, adjacent spots can be spatially overlapped (Sun et al., 2018).

Both of these studies use Gaussian beams, which is technically simpler compared to the 3D-CGH-TF system proposed by (Hernandez et al., 2016). However, a considerable drawback of this system is the fact that the laser beam is focused on a line on the SLM used for 3D multiplexing, where the size depends on the linear dispersion of the TF system in the direction of dispersing the spectral frequencies (usually parallel to the optical table), and it equals the monochromatic beam size in the unchirped dimension (vertical to the dispersion direction). In the chirped direction, care is usually given to fill the size of the SLM liquid crystal array, while in the unchirped direction the size of the beam is a few millimeters (2–3 mm) (Durst et al., 2008; Sun et al., 2018). This imposes a restriction to the maximum laser power used before damaging the SLM (and thus to the maximum number of spots that can be projected). To overcome this limitation Pégard et al. (2017) proposed adding an extra lens before dispersing the ultrashort pulses on the TF grating, defocusing the beam on the SLM and thus increasing the illumination area. Nevertheless, the defocusing created a secondary spatial focus in the form of a line that deteriorated the axial propagation of the beam (Pégard et al., 2017; Chen et al., 2018c). In a more advanced version of this method the lens was replaced by a rotating diffuser at an image plane, after the grating. This led to an enlarged illumination of the SLM, without secondary focus effects (Mardinly et al., 2018) and Gaussian beams of 23 μm axial FWHM (axial optical Point Spread Function; PSF) (Pégard et al., 2018). 3D-SHOT was used in this case to simultaneously stimulate neurons co-expressing newly developed excitatory or inhibitory somatic opsins, ST-ChroME or IRES-ST-eGTACR1, respectively, and GCaMP6s in an all-optical configuration (about 50 neurons in three different planes extending to an axial range of 100 μm, 0.13 mW/μm² or 40 mW per target for activating ST-ChroME neurons, 0.08 mW/μm² or ∼6 mW per target for IRES-ST-eGTACR1 neurons, illumination with a low-repetition rate laser at 1040 nm for 1 s) (Mardinly et al., 2018).

Higher flexibility and better axial resolution was demonstrated by Accanto et al. who presented a system for multiplexed temporally focused-light shaping (MTF-light shaping), where the beam-shaping part was either 2D-CGH or GPC (Accanto et al., 2018) (Figure 2C,D). For MTF-CGH, the optical setup is the same as the one of 3D-CGH-TF (Hernandez et al., 2016), the only difference being the way the multiplexing SLM is addressed. MTF-CGH enabled investigators to generate 15-μm diameter temporally focused holographic spots on 50 independent planes on an excitation field of 300 × 300 × 500 μm³ and an average axial PSF of 11 μm FWHM. The theoretical FOE for the optical parameters they used was 750 × 750 × 990 μm³, but was experimentally limited by the size of the optics used.

In MTF-GPC the beam-shaping part was substituted by a GPC setup (Glückstad, 1996; Papagiakoumou et al., 2010) for projection of high-precision, speckle-less, temporally focused arbitrarily shaped patterns. Notably, the combination of GPC with CGH to extend GPC to 3D was previously reported (Go et al., 2011; Bañas and Glückstad, 2017) but without TF, which is essential for suppressing excitation by the out-of-focus light (Papagiakoumou et al., 2010). In MTF-GPC, characterization of 12-μm diameter TF-GPC spots showed improved axial PSF, compared to MTF-CGH (6 μm FWHM on average on a FOE of 200 × 200 × 200 μm³), as expected for TF-GPC patterns (Papagiakoumou et al., 2010). Similar to the Gaussian beam case, a crucial drawback remains the illumination of the multiplexing SLM with a line. To overcome this, (Accanto et al., 2018) removed the phase contrast filter and used the first SLM of their configuration to perform amplitude and phase modulation. They encoded a pattern in four different areas on the beam-shaping SLM and the pattern of each area was laterally displaced with a different prism-phase effect, such as generating four different lines on the multiplexing SLM after the beams were temporally focused. Addressing the first SLM in different areas enabled the projection of replicas of four different speckle-free patterns in a volume of 300 × 300 × 400 μm³, a method referred to as MTF-Multi Shapes. This strategy both increased the illumination area of the multiplexing SLM and allowed more flexibility on the shape of the projected patterns, similar to that described by (Hernandez et al., 2016).

Evidently, flexibility of the MTF-light shaping methods comes at the cost of simplicity of the optical setup and total cost. For simpler cost-effective solutions in applications where the excitation spot size and form can be predetermined, it is possible to use static lithographically fabricated phase masks (Accanto et al., 2018) to replace the first SLM in MTF-CGH, or a GPC-light shaper (Bañas et al., 2014) to replace the GPC setup in MTF-GPC.

MTF-CGH was used in a multi-cell excitation of photoactivatable GCaMP (paGCaMP) in the central nervous system of the drosophila larvae (photoconversion at 760 nm with 1.0 mW/μm², illumination with trains of 10-ms pulses at 50 Hz, total illumination duration: tens of seconds up to 4 min) and to photoconvert Kaede in the zebrafish larva hindbrain (photoconversion at 800 nm with 0.4 mW/μm², illumination with trains of 10-ms pulses at 50 Hz, total illumination duration 1–4 min). Parallel illumination of neurons allowed fast photoconversion in both cases, with minimal photoactivation of untargeted neighboring cells. Especially in the case of paGCaMP, neuronal processes of the targeted cells could be clearly distinguished from the background, allowing the possibility to precisely track neuronal morphology (Accanto et al., 2018).

Although the use of TF in neuronal photoactivation with parallel methods has offered the possibility to locally confine the excitation volume, such as to preserve single-cell resolution, this might not be necessary for low-scattering samples, excitation of small size cells or sparse staining. In that case, 3D-CGH alone can be used for the projection of extended light patterns in different planes (Haist et al., 1997; Hernandez et al., 2016). Thus 2P 3D-CGH spots of 6-μm diameter were used to photoactivate ChR2 in the zebrafish larval neurons, which in combination with 2P GCaMP6s Ca²⁺ imaging, enabled the identification of neuronal ensembles associated with control of tail bending (photoactivation at 920 nm, 0.2 mW/μm² or 50 mW per target, circular overlapping photoactivation regions of 18 μm in diameter) (Dal Maschio et al., 2017). Moreover, the authors performed targeted photoactivatable GFP (paGFP) photoconversion to obtain a morphological reconstruction of individual functionally identified neurons (photoconversion at 750 nm with 0.25 mW/μm² or 7 mW per target, 1-s illumination).

Finally, as a general comment we should note that 3D-TF methods using CGH for multiplexing the excitation spot, can produce powerful experimental configurations in volumetric FOEs, with the quality of all spots being the same across the whole FOE, since the spots multiplexed are replicas of the same single original spot. However, for volumes reaching mm-range, care should be taken to homogenize the excitation properties of the projected spots, by taking into account factors such as scattering with increasing depth, SLM-diffraction-efficiency corrections, optical aberrations due to the large defocusing of spots (objective used at its limits) or projection near the borders of the FOE, and spectral aberrations for TF occurring by cropping spectral frequencies in the optics when using large defocus (Hernandez et al., 2016).

Scanning methods are alternative approaches for neuronal stimulation and have been widely used in 2P optogenetics (Rickgauer and Tank, 2009; Andrasfalvy et al., 2010; Packer et al., 2012; Prakash et al., 2012), mainly for the activation of the slow kinetics opsin, C1V1, and most of the times in a spiral trajectory (Packer et al., 2012, 2015; Carrillo-reid et al., 2016; Yang et al., 2018). They represent the simplest and most immediate solution for many laboratories, since they adopt conventional 2P scanning microscopes based on galvanometric scanners. Nevertheless, the sequential photostimulation limits the achievable temporal resolution (Papagiakoumou, 2013; Ronzitti et al., 2017b) and does not allow the simultaneous activation of multiple targets. The use of resonant scanners or acousto-optic deflectors (AODs) to increase the temporal resolution of scanning methods, is still limited by the necessary dwell time in 2P excitation, especially for slow opsins (Prakash et al., 2012). Moreover, due to their cycling at resonant frequencies (8 kHz), resonant scanners cannot provide the necessary flexibility for arbitrary excitation trajectories, like spiral scans. For simultaneous multicell activation in 3Ds, scanning microscopes can be modified to include a SLM to multiplex the beam prior to the scan via 3D-CGH (Packer et al., 2012; Yang et al., 2018). The holographic pattern, consisting of multiple near-diffraction limited spots (∼1 μm in diameter), is then scanned in spiral trajectories on an area covering the size of the cell soma (Rickgauer and Tank, 2009; Packer et al., 2015; Yang et al., 2018).

Studies using hybrid methods in all-optical configurations have so far used C1V1 as optogenetic actuator excited at 1040 nm (Packer et al., 2015; 20–80 mW per target, spirals of 20 ms, 80 MHz repetition rate laser; Yang et al., 2018; 2.2–6.0 mW per target, spirals of 20 ms, low-repetition rate amplified laser). For 3D manipulation of neurons specifically, there is currently only one study where investigators simultaneously photostimulated more than 80 neurons over 150 μm in depth in layer 2/3 of the mouse visual cortex, while simultaneously imaging the activity of the surrounding neurons with GCaMP6s (Yang et al., 2018). The authors photoactivated in three different planes in an axial range of 100 μm selected groups of somatostatin inhibitory interneurons, suppressing the response of nearby pyramidal neurons to visual stimuli in awake animals (6 mW per cell from a low-repetition rate laser, or 6 mW/μm², since the surface of the illumination spot in that case is about 1 μm², illumination for 2.8 s with 175 continuously repeated spirals, each lasting ∼16 ms).

An estimate of the maximum possible number of targets to address with each approach in the framework of an all-optical experiment, presumes knowledge of the total light losses of an optical system from the laser source to the objective output, which can significantly vary from one system to another. Moreover, the power necessary per target used, can vary according to opsin type, expression level, cell health and activation depth. In general, power losses for parallel illumination methods mainly consist of losses on the SLM(s) and the diffraction grating used for TF. For hybrid methods losses are approximately 2–3 times less than those of parallel ones, since they are mainly due to the use of a single SLM. It has also been reported that parallel approaches need about twice the power used by a spiral scanned beam, to induce a neuronal response with the same properties in both cases (Picot et al., 2018; Yang et al., 2018). Thus, in principle, for the same laser source and systems carefully designed to minimize power losses, hybrid methods are supposed to outmatch parallel ones about 4 times in the maximum number of cells possibly activated (without considering photodamage limits). From what is reported so far in literature, the maximum number of cells that have been simultaneously activated with parallel approaches is 50 neurons (Mardinly et al., 2018), while for hybrid methods this number is reported to be up to 80 (Yang et al., 2018). In the first case, the authors clearly state that they were limited by the available power of their laser system. However, current advances in laser technology can provide fiber amplified systems that deliver up to 60 W maximum average power, allowing for the possibility to greatly increase the above reported numbers.

A fundamental difference between parallel and hybrid 3D multi-target photoactivation methods is the eventual photodamage effects that one can induce by increasing the number of targets and thus the amount of light that is sent to the tissue. As presented in the previously reported cases, in general, parallel illumination approaches use lower illumination intensity [<0.4 mW/μm² independently on the opsin type, see also (Chen et al., 2018b)] but higher total average power per target (e.g., for 2P in vivo activation 10–45 mW when amplified laser pulses of tens of μJ pulse energy are used from low-repetition rate fiber amplifiers, 30–90 mW when nJ-energy pulses are used from MHz-repetition rate oscillators) than scanning approaches (Chen et al., 2018b; Forli et al., 2018; Mardinly et al., 2018), making them more vulnerable to heating, i.e., linear photodamage. On the other hand, scanning (or hybrid) methods use high intensity (2–6 mW/μm²) focused beams of low total average power (Yang et al., 2018) (although average powers in the range of 20–80 mW were reported (Packer et al., 2015) when nJ-energy pulses were used from MHz-repetition rate oscillators), making them vulnerable to non-linear photodamage.

For non-linear photodamage, the damage threshold was shown to be inversely proportional to the pulse duration and proportional to the square of the mean power (König et al., 1999). It has been evaluated on the basis of morphological damage for CHO (Chinese Hamster Ovarian) cells to 0.1 J/cm² (König et al., 1999), or tissue ablation for porcine cornea (Olivié et al., 2008) to 1.5–2.2 J/cm² for 800–1000 nm. For comparison, an intensity of 80 mW/μm² at 80 MHz pulse repetition rate corresponds to a fluence of 0.1 J/cm². No relevant studies exist for the mouse brain. A recent study on tissue heating, took the standard illumination parameters for either parallel or scanning methods into consideration, showing that the local temperature rise on a target area did not exceed the physiological limits in both cases (<1 K) (Picot et al., 2018). Specifically, to generate an action potential in vivo, with a holographic spot of 12-μm diameter at a depth of ∼150 μm, illuminating a neuron for 3 ms, at 1030 nm and 0.1 mW/μm², the average temperature rise over the spot’s surface, is estimated to 0.3 K. Furthermore, comparing the temperature rise for experimental conditions able to generate action potentials with latency in the range of 2–10 ms in vitro, it was found that for a holographic spot (3-ms illumination, 0.2 mW/μm²) the average rise was 1 K, while for a focused beam in a spiral trajectory (3-ms illumination, 31 mW/μm²) the mean temperature rise was <0.5 K, and the local rise at the center of the spiral was again ∼1 K (Picot et al., 2018). For multi-target excitation, what remains critical is the distance between the different targets: for spots placed at an average distance from their nearest neighbor greater than the thermal diffusion length in tissue lth=6Dt, where D is the thermal diffusion constant [140 μm²/ms (Yizhar et al., 2011), and t the evolved time], the temperature rise remains comparable to the case of the isolated spot (for 3-ms illumination duration, holographic stimulation at an intensity ≤0.2 mW/μm² enables keeping the induced temperature rise <2 K for activating 100 cells whose inter-soma distance was larger than the thermal diffusion length, which in that case was ∼50 μm). Otherwise, the heat load starts to significantly increase locally (Picot et al., 2018). Moreover, in terms of illumination duration, prolonged stimulation (>1 min) was found to induce substantial brain heating (6–8°C) (Mardinly et al., 2018).

Notably, the considerations of the study by Picot et al. (2018) indicate that the optimal laser repetition rate for 2P optogenetics depends on the adopted illumination method: the very low excitation intensity used in parallel illumination, allows to neglect non-linear damage effects and privileges using low (500 kHz–2 MHz) repetition rate lasers to minimize heating through linear absorption. Scanning approaches on the other hand, require high excitation intensity but enable more efficient heat dissipation because of their small beam spot size; therefore, for short illumination times, a higher repetition rate laser (Ji et al., 2008) should be preferred in order to minimize peak-power-sensitive damages.