Reshaping the optical dimension in optogenetics

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Reshaping the optical dimension in optogenetics Alipasha Vaziri1 and Valentina Emiliani2 Optogenetics has been revolutionizing circuit neuroscience in the last few years. Optical methods combined with genetics and molecular techniques have provided new tools for stimulation of neurons, which hold great promise to provide a solution to the circuit mapping problem and more generally provide us with the ability to artificially control the natural stimulus space. Nevertheless, until very recently almost all applications of optogenetics have been based on relatively simple optical schemes mainly used for inducing population activity in neuronal assembles. In this context, alternative optical schemes that enhance the spatial or temporal resolution of excitation and allow for flexible and arbitrary generation of light patterns have all synergetic impact on the development of new optogenetic actuators. In the following we discuss and compare the main new optical techniques that have become available in the recent years. Their respective strengths and limitations as well as their application to different biological contexts are illustrated. Addresses 1 Research Institute of Molecular Pathology (IMP) & Max F. Perutz Laboratories, University of Vienna, Austria, Dr. Bohr Gasse 7-9, A-1030 Vienna, Austria 2 Wavefront-Engineering Microscopy Group, Neurophysiology and New Microscopies Laboratory, Centre National de la Recherche Scientifique, Unite´ Mixte de Recherche 8154, Institut National de la Sante´ et de la Recherche Me´dicale U603, Paris Descartes University, Paris, France Corresponding authors: Vaziri, Alipasha ([email protected]) and Emiliani, Valentina ([email protected])

Current Opinion in Neurobiology 2012, 22:128–137 This review comes from a themed issue on Neurotechnology Edited by Winfried Denk and Gero Miesenbo¨ck Available online 29th December 2011 0959-4388/$ – see front matter Published by Elsevier Ltd. DOI 10.1016/j.conb.2011.11.011

Introduction A wide range of brain functions, such as the mechanisms underlying the storage of memories or the generation of motor behavior rely on the dynamic interaction of sensory inputs with neuronal circuits. Understanding the neuronal basis for such complex phenomena requires solving one of the major challenges in current neuroscience: the wiring diagram of the neuronal circuits. In addition one has to be able to monitor and control parallel computations occurring at multiple levels: from sparse coding of information by neuronal ensembles to events Current Opinion in Neurobiology 2012, 22:128–137

occurring subcellularly at the level of individual dendrites and synapses of neurons. Yet, given the complexity of neuronal circuits this task to a great extent has been hindered by the lack of appropriate tools. In the recent years, the combination of optical methods with genetic techniques has led to significant advances to address both problems: the neuronal stimulation and recording. For the purpose of this review, we will focus on methods based on optogenetics [1,2,3,4], which have in many cases provided a solution to the stimulation problem. In these techniques the genetic expression of light gated channels and pumps such as channelrhodopsin (ChR2) [5–7], light gated glutamate receptors (LiGluR) [8] and halorhodopsin [9] that can be activated by visible light illumination has enabled the optical initiation [10,11,12] and inhibition [13] of population activity in defined neuronal populations and optical control of behavior [9,14–16]. Interestingly, up to now the low excitation level (few mW/mm2) required for the optical activation of optogenetics actuators combined with the specificity of genetic targeting have permitted solving key biological questions with relatively simple illumination methods based on wide field visible light illumination. Nevertheless, until very recently it has been proven challenging to optically control a specific subset of neurons or neuronal process in a targeted fashion in a densely packed three dimensional neuronal population of genetically identical neurons. For this, more flexible optical approaches are needed permitting fast, precise and in-depth light delivering at specific regions. The availability of such a tool combined with calcium imaging would for instance allow mapping neuronal circuits in a high throughput fashion. Also by mimicking the natural stimulus space of the sensory neurons and controlling the spatiotemporal dimensions of this space one would eventually be able to move from a correlative to a causal relationship between stimulus, network activity and behavior. In the following we discuss and compare a few optical techniques that have recently been experimentally demonstrated to achieve this task and bear the potential to enter the tool-kit of the broader neuroscience community. The general idea in these techniques is to control the spatial excitation volume by patterning or ‘sculpting’ the laser beam while minimizing the time to induce, for example, an action potential or a Ca2+ response and maintaining flexibility and speed in generation of arbitrary sequences of spatio-temporal excitations patterns. Based on the light source used and the molecular excitation mechanism, they fall into two broad categories of single photon and two photon patterned stimulation. www.sciencedirect.com

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of identical neurons [10,17] a single cell [17,18– 20,21,22] or a single cell process [23,24,25] has been demonstrated using two different approaches; amplitude and phase modulation.

Single photon patterned stimulation In single photon excitation, the absorption of a single photon leads to the activation of the photosensitive actuator. Single photon patterned activation of a subset

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Single photon light patterning. (a) Left, schematic of an amplitude modulation beam shaping with a digital light projector (DLP)-based photostimulation setup. Top right, dorsal surface of the olfactory bulb with a tetrode positioned in the mitral cell layer. One square light spot can be seen projected onto the bulb surface. Scale bar represents 500 mm. Inset, cartoon schematic of glomeruli showing a subglomerular size light spot and dual-tetrodes positioned in the mitral cell layer. i, DLP projector; ii, focusing lens; iii, blue excitation filter; iv, dichroic mirror; v, emission filter; vi, CCD camera; vii, dual-tetrode; viii, olfactory bulb. (b) Top, raw voltage traces corresponding to the four channels of a tetrode during photostimulation. Center, raster plot of spikes from an isolated single unit. Bottom, peri-stimulus time histogram (PSTH) with 20-ms time bins. Adapted from [22]. (c) Schematic of a holographic light patterning set up. A visible cw laser beam (i) is sent after expansion (ii) on a LC-SLM (iii), a first lens (iv) generates the image of the illumination pattern at its focal plane, the image is rescaled at the sample plane via a telescope formed by a second lens (v) and the microscope objective (vi); vii, CCD camera. (d) Top left, wide-field epi-fluorescence image of a CA1 hippocampal pyramidal cell and selected region of interest (white box) used to obtain the phase-hologram and the corresponding Illumination pattern at the objective focal plane (top right). Bottom left, wide-field fluorescence image of CA1 hippocampal neurons loaded with oregon green bapta and selected region of interest (white line) used to obtain the Illumination pattern at the objective focal plane (bottom right). Adapted from [17,23]. (d) Phase profile at the back aperture of the objective used to generate a circular spot of 10 mm in diameter and corresponding axial propagation for a Gaussian (left) and a holographic beam (right). In the case of non-diffraction limited spot, DH permits improved axial resolution. The underlying reason for this effect is the spatial shape of the optical wave front or phase profile at the back aperture of the objective. In contrast to slowly varying spatial phase of a Gaussian beam (left) when light patterning is generated via DH (right) the phase profile is highly inhomogeneous, the presence of high frequency component permits illuminating the objective with higher angles so that final patterns are generated without the need of scarifying the effective NA of the beam. For example, the holographic beam in the right shaped to selective excite an area of 10 mm in diameter corresponds to an axial resolution of about 20 mm [17,23,34]. Comparing this number to the value of >1 mm achievable if the same area is covered by under filling the objective rear aperture with a Gaussian beam (left) or by using DLP patterned light, it represents a significant improvement in the axial resolution. Adapted from [23]. www.sciencedirect.com

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Different schemes for high resolution two-photon optogenetics. Spiral scanning. (a) Two-photon activation of channelrhodopsin expressing neurons requires the recruitment of a large number of channels that generate sufficient inward current to induce an action potential within a time window shorter than the channel deactivation time. In Spiral scanning this was achieved by utilizing a spiral path (left) and increasing the area of the excitation spot by underfilling the back aperture of the objective (i.e. by reducing the effective NA) (center). While the spiral scanning path reduces the required scanning time compared to the raster scanning, the larger excitation spots leads to fewer number of positions that have to be visited. In this scheme action potentials on single neurons were achieved within a 32 ms spiral scan and an effective NA of 0.3. lexc = 920 nm. Adapted from [43]. (b) Serial Light sculpting. Left, the serial light sculpting relies on the integration of temporal focusing (TF) into a galvanometer laser scanning head [43]. A temporally focused spot is generated at the intermediate plane between the scan and the tube lens of a two-photon scanning microscope which is then imaged into the sample plane via a telescope consisting of the tube lens and the objective. In this scheme the fast stimulation of a single neuron is achieved by the simultaneous activation of a large number of channels via temporal focusing while targeted high speed spatiotemporal patterning of a population with single cell resolution is achieved by rapid repositioning of the sculpted beam with the scanning mirrors. This has been demonstrated on CA1 pyramidal cells [46] and interneurons in acute hippocampal slices (center, the figure on top shows the targeted stimulation of somata of a set of 25 interneurons in a hippocampal slice (gray spots) and the postsynaptic responses recorded). By using different spatiotemporal sequences of excitation different temporal patterns of postsynaptic excitation or inhibition can be generated on a target neuron. One application of this has been the modulation of the shape of the perisomatic inhibition [46]. The different spatiotemporal patterns of excitation have been used to mimic a symmetric or delayed perisomatic inhibition (right) and test different hypothesis about the underlying mechanism of theta related neuronal activity; lexc = 880 nm, 300 mW out of the microscope objective, spot size 7 mm, 120 fs pulses. Adapted from [46]. (c) Parallel Light sculpting. Left, layout of TF-GPC set up: Output of Ti:Sapphire laser is expanded to illuminate the LC-SLM that is located at the front focal plane of a 4-f imaging setup (L1, L2), with a phase contrast filter (PCF) at the confocal plane between the lenses. This generates at the grating plane the desired lateral patterning. For the control of the axial resolution through temporal focusing, a blazed reflective grating is placed at the front focal plane of a 4-f imaging setup comprising the lens L and the objective. Center, shaped pattern (bottom) based on a confocal image of a Purkinje cell in selected regions of interest (yellow lines) (top). Right, wide-field epi-fluorescence image of cell body loaded through the patch pipette with Alexa 594 and AP trains evoked following light stimulation at

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In amplitude modulation, patterned stimulation is achieved by using a micro LED array [19,21] or a digital light processing (DLP) mirror array [10,18,20,22,26] (Figure 1a, b). In the latter, orienting the mirrors in the off position permits redirecting the light out of the target area thus generating the desired pattern [27]. Phase modulation approaches make use of liquid crystal spatial light modulators (LC-SLM) [17,23,24,25,28,29,30]. In this case the excitation shape is generated by using the principle of digital holography (DH). Originally proposed for generating multiple optical tweezers [31,32], the experimental scheme for DH consists of computing the phase pattern at the rear aperture of the objective that reproduces the desired target intensity in the objective focal plane using an iterative algorithm. The calculated phase-hologram is imposed onto the input beam wavefront by a LC-SLM. In this way, after propagation through the objective the desired pattern is produced at the focal plane. The use of DH allows for a high level of flexibility in the sense of permitting simultaneous excitation of different patterns ranging from single to multiple diffraction limited [23,24,25,29] spots, to shapes covering a sub cellular region [17,23] or multiple sparse cell [17] and spots arranged in 3D configurations [24,25,33] (Figure 1c). In the case of a non-diffraction limited spot, DH permits also improved axial resolution (Figure 1d) [17,23,34]. However, in DH the spatially varying diffraction efficiency of liquid crystals (see below) and the presence of spurious light (zero order spot and ghost images) [35] limits the maximum size of the excitation field achievable [17,29]. Also, one intrinsic limitation in the use of DH, not present with amplitude modulation techniques, is the presence of intensity fluctuations (speckles) inherent to the approximations in the holographic iterative algorithm. Solutions based on using a rotating diffuser [36] or phase mask shiftaveraging [37] permit averaging over the speckles and lead to a smoother spatial distribution, however at the expense of temporal resolution and light efficiency. With phase or amplitude modulation light patterning, it has been possible to activate the left or right side of spinal cord segment in LiGluR-expressing zebrafish larvae and to correlate the spatial localization of stimuli with directional tail movements [10]. Further applications in optogenetics include in vivo stimulation of individual glomeruli expressing ChR2 for identifying mitral/tufted cells receiving common inputs [22] and patterned activation of specific neurons expressing ChR2 combined with G-Camp fluorescence imaging for functional mapping of neuronal circuits in C. elegans [20]. The use of single photon holographic photoactivation has been lim-

ited so far to the neuronal activation through glutamate uncaging [17,23,24,25] but its extension to optogenetics will be straightforward. The underlying physical principles of single photon excitation leaves these techniques with two fundamental limitations; low optical sectioning and strong scattering of visible light in brain tissue, so that 1P patterned experiments have been limited in brain tissue to depths in the range of 20–40 mm [17,22,23,24].

Two-photon patterned stimulation The use of two photon excitation [38–40] is the obvious choice for improving penetration depth and axial resolution. In this case, excitation is achieved via a simultaneous absorption of two photons with each having half of the energy of the photon in single photon excitation. The conditional probability of this process results in the intensity squared dependency of the molecular excitation and the use of longer wavelength permits increasing the penetration depth. Two-photon excitation has been widely used to achieve high optical sectioning at depth and to reduce photo-bleaching in imaging. However, exploiting the advantages of two-photon microscopy for optogenetics is more involved. One issue has been the low (approximately 40 fS) [12,41] channel conductance of ChR2 and in general of optogenetic channels and pumps and the fast deactivation time. As a result, the typical excitation volume (1 mm3) used in two-photon scanning microscopy is not suited for recruiting sufficiently large number of channels within a short time period to, for example, optically induce an action potential. To achieve this, a large number of channels covering an area about the size of the soma need to be activated near simultaneously while axially confining the activation light to the size of the cell body or of a single cell process. In the last years different optical schemes have been proposed [36] and shown to achieve this task [42,43,48]. In the first demonstration of two-photon activation of ChR2, a laser scanning approach [42] was used that allowed for a sufficient integration of individual channel currents for inducing an action potential. Combing a low numerical aperture (NA) beam adjusted to match the axial dimension of the cell and a spiral shaped path on the soma for faster scanning, the channel deactivation could be outran and the initiation of an action potential could be shown within a 30 ms time window (Figure 2a). The temporal resolution of this approach can in principle be enhanced by further underfilling the back aperture of the objective (i.e. reducing the effective NA) which would lead to a larger spot on the soma and an increase of the number of channels that are excited simultaneously. However, given the inverse quadratic dependence of the axial

10 Hz, 20 Hz, and 30 Hz (0.40 mW mm2; 10 ms pulse; 15 mm excitation spot). Scale bars: 20 mm. lexc = 920 nm, 40, 0.8 NA objective, excitation depth of 50–70 mm (top). Transmission image of two ChR2-YFP positive neurons (A, B) with 15 mm excitation spots (red) superimposed. Neurons A and B were simultaneously patch clamped. A spot over neuron A triggered an AP only in A. A spot over neuron B triggered an AP only in neuron B. When the two cells were illuminated simultaneously both A and B fired an AP (0.25 mW mm2; Scale bars 10 mm; lexc = 850 nm, 40, 0.8 NA objective). Recordings in 10 mM NBQX (bottom). Adapted from [48]. www.sciencedirect.com

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Box 1 Spatial and temporal focusing The signal generated by two-photon excitation is given by: S2P / P2/tf, where P is the average power density, t the pulse duration and f the laser repetition rate. In the conventional scheme for two-photon scanning microscopy, the pulse duration is constant during the axial propagation and owing to the non linear dependence of the two-photon signal, S2P, on P, axial localization of excitation is achieved by spatially focusing the excitation beam on a diffraction limited spot by overfilling the back aperture of the objective (a). Under these conditions, the axial (zR) and the lateral (w0) beam parameters are coupled via a square law: 2zR = pw02/l / l/NA2, where NA is the objective numerical aperture. In temporal focusing a comparable axial resolution can be achieved for non-focused (low NA beam). This is obtained by using the frequency spectrum of the pulse to axially control, via dispersion, the pulse duration, t. The pulse is short at the focal plane and becomes longer in the out of focus regions (b). In this fashion, utilizing the fact that the fluorescence signal (S2P) is inversely proportional to the pulse duration (t) axial localization of excitation is achieved by temporally focusing the excitation beam. Under these conditions it can be shown [50,55] that the axial resolution is independent on the excitation spot size and an effective decoupling of the axial and lateral beam parameters is reached.

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resolution, 2zR, on the NA (2zR / l/NA2) this would correspond to a dramatic loss of axial resolution. Hence, in order to achieve high temporal resolution and axial confinement to a single cell a decoupling of the axial and lateral beam parameters is required. Andrasfalvy et al. [43] have recently demonstrated such a decoupling by using the technique of temporal focusing (TF) [44,45] (Box 1). ‘Sculpted’ light distributions that are axially and laterally confined to the target neurons (Figure 2b) were generated and used for inducing single cell firing with 2 ms temporal resolution and 40 mm axial resolution when dendrites as axially confined regions of neurons where used for measurements [43].

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circular excitation areas and requires readjusting the spot size with a variable telescope for experiments requiring different excitation spot sizes. To generate arbitrary spatiotemporal excitation patterns that match the shape of neuronal processes or of targeted neuronal populations, this technique was integrated into a laser scanning head. Using such sculpted light patterns that were serially applied in controlled spatiotemporal sequences, action potentials from a single dendrite in acute brain slices could be triggered at 150 mm depth by placing multiple spots along the dendrite [43] with 6 ms temporal resolution. The same approach has allowed to study the underlying mechanisms of theta phase precession by inducing spatiotemporal patterns that mimic the perisomatic inhibition on hippocampal pyramidal cells (Figure 2b) [46].

In this case by saturated excitation of ChR2 the axial resolution was sacrificed to achieve high temporal resolution for inducing an action potential. In other applications where only subthreshold neuronal stimulation is desired the excitation power can be reduced which would lead to an enhancement of the above values.

Although the combination of ‘sculpted’ light with fast laser scanning permits improving the temporal resolution compared to the spiral scanning approach without deterioration of the axial resolution, still it does not permit truly simultaneous excitation of multiple locations.

While this approach allows a simple implementation of decoupling of the lateral and axial resolution, it is limited to

This limitation can be overcome by ‘sculpting light’ with a full parallel approach. For parallel light patterning,

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Optical dimension in optogenetics Vaziri and Emiliani 133

methods based on amplitude modulation are far too inefficient for 2P excitation, so the only option is to use phase modulation. This has been achieved first by DH [33,36,47] and later [48] with the generalized phase contrast method (GPC) [49]. Briefly, with GPC the target intensity map is converted into a binary phase map which is transferred to the input beam wavefront via the LC-SLM. The modulated beam is then focused on a phase contrast filter that imposes appropriate phase retardation between the on-axis focalized component and the higher-order Fourier components. The interference between these two beams generates the original target intensity at the focal plane of a second lens (conjugate with the objective focal plane). Combining GPC and temporal focusing (Figure 2c), the generation of parallel arbitrary excitation patterns attaining the limiting axial resolution of temporal focusing has been demonstrated and used for efficient in-depth two photon-activation of ChR2, enabling for the first time, the simultaneous excitation of multiple neurons or multiple neuronal compartments in cultured neurons and brain slices [48].

Discussion Overall each of the above approaches has certain advantages and drawbacks and, as a result, each of them might prove to be best suited in a given experimental setting as summarized in Table 1 (see also Ref. [50]). One major difference between the scanning and parallel approaches based on phase modulation is the higher required average laser power in the latter. In digital holography, simultaneous stimulation of multiple areas implies that the available laser power is dived by the total excitation area, this combined with the intensity squared dependency of excitation in the two-photon process means that, at a given laser power, the induced inward current will decrease non-linearly with the total area excited and could lead to some practical constrains on the total number (and area) of the spots that can be excited simultaneously and the maximum penetration depth in brain tissue. Moreover, the pixelated structure of the LCSLMs limits the conversion efficiency (available excitation power/output laser power) to 70% (the rest goes in the zero order and ghost images) and determines a position (x,y) dependent variation of the spot intensity inside the excitation field which decreases to zero proportionally to (sin(x)/x)2(sin(y)/y)2 [17,29] and to (sin(x)/ x)4(sin(y)/y)4 for 2P excitation [48]. This sets also a limit on the size of the maximum excitation to few hundreds of mm2 [17,24,29].1 Compared to DH, the advantage of 1 the maximum excitation field is given by  In digital holography,   f 1 f obj , where f 1/f 2 is the telescope magnidxmax ¼ dymax ¼ 2  2dlSLM f2 fication (typically 1.5), f obj is the objective focal lens and 1/2dSLM is the SLM spatial frequency (10 lp/mm for 50% efficiency in 2P). This for l = 900 nm leads to an effective excitation field of 120 mm  120 mm for a 40 objective.

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GPC is that the generation of light patterns does not require the use of iterative algorithms thus permitting the generation of speckle free patterns. Moreover, for GPC the theoretical conversion efficiency can be 100% providing that the total excited area reaches 1/4 of the total excitation field. However, this sets a limit on the maximum excitation density achievable in GPC which is given by the ratio between the available excitation power and the excitation field, regardless of the excitation spot size. This has in practical applications limited the excitation field to a circle of 60–100 mm in diameter [48]. By contrast, serial light sculpting where all the available power can be concentrated in a single spot scanned trough the sample, it has less stringent power requirements permitting a greater depth penetration in brain tissue at a given laser power. Further, the use of scanning mirrors allows for a larger excitation field and more uniform light distribution on a larger excitation area (0.5 mm  0.5 mm for a 40 objective). Another key point in the comparison between these two approaches is the temporal resolution. In order to include both the necessary time to generate a single excitation pattern, k (e.g. induce a certain amount of inward current) and the required time to apply a specific sequences of, n, spatial patterns, the temporal resolution Ts can be defined as Ts ¼

n X

N k  ðRt þ S t Þk þ ðk  1Þ  1=Re ;

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Here, N is the number of visited positions to cover one pattern and is given by the ratio between the total excited area of the pattern and the excitation spot size, Rt the residence time at a given position (also known as dwell time), that is, for optogenetics the time to induce a certain amount of depolarization – that is to induce an action potential or to induce a subthreshold stimulation or a detectable Ca2+ response –, St the point to point scanning time and Re the LC-SLM refreshing rate (60– 200 Hz) – that is the inverse of the time needed to address the LC matrix with a new patterns without having memory of the previous one. In parallel approaches, N = 1, St = 0, and the temporal resolution, Ts, for a single excitation pattern (n = 1) is only limited by the residence time Rt. For scanning approaches Ts = N  (Rt + St) and a compromise between the choice of N and Rt, and the achievable axial and temporal resolution has to be found. First, the overall Ts can be optimized by reducing St, as it was done in the spiral scanning approach [42]. Moreover, the number of visited positions N can be decreased by increasing the excitation spot area, that is, by reducing the beam numerical aperture, however this leads to a quick loss in axial resolution. In serial light sculpting, where a decoupling of lateral and axial Current Opinion in Neurobiology 2012, 22:128–137

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Table 1 Overview of the main patterned light neuronal stimulation approaches

Spiral Scanning

Serial Light Sculpting

Parallel Light sculpting (DH or GPC)

Optical set up:

Optical set up:

Optical set up:

Conventional 2P scanning microscope

Conventional 2P scanning microscope + TF

Conventional 2P scanning microscope + TF + DH or GPC

Maximum reachable excitation density:

Maximum reachable excitation density:

Maximum reachable excitation density:

= Laser power divided by the excitation spot size

= Laser power divided by the excitation spot size

Power losses:

Power losses: • ~15-20% loss at the grating

(DH)= Laser power divided by the total excitation area

(GPC)= Four times the laser power divided by the excitation field

Power losses: • ~15-20% loss at the grating

• ~15-20% loss at the grating

• ~30% loss at the SLM Temporal resolution for single pattern:

Temporal resolution for single pattern:

Temporal resolution for single pattern:

= Residence time plus point to point scanning time, multiplied by the number of excited position

= Residence time plus point to point scanning time, multiplied by the number of excited position

= Residence time

Temporal resolution for sequential patterns:

Temporal resolution for sequential patterns:

Temporal resolution for sequential patterns:

= Number of patterns multiplied by the temporal resolution of a single pattern

= Number of patterns multiplied by the temporal resolution of a single pattern

≅ Number of patterns

Axial resolution: Proportional to the square of the lateral excitation spot size

Axial resolution: Independent of the lateral excitation spot size

Axial resolution: Independent of the lateral excitation spot size

Excitation field: limited by galvanometric scanners (~0.5x0.5 mm 2 for 40x objective)

Excitation field: limited by galvanometric scanners (~0.5x0.5 mm 2 for 40x objective)

Excitation field: • (DH) limited by LCSM diffraction efficiency (few hundreds of μm 2 )

resolution is achieved this constraint is removed and, in the limit of single cell excitation (N = 1, St = 0), Ts is only limited by the residence time, Rt. For patterns requiring a large number of N, Ts can be enhanced by reducing Rt. This can be achieved by increasing the excitation power [42,43], however at the cost of lateral and axial resolution owing to the strong contribution of the out-of-focus light to the evoked responses [42,43,46]. Current Opinion in Neurobiology 2012, 22:128–137

multiplied by the LC refreshing rate

• (GPC) given by the parameters of the interferometer; practically limited by power requirement; typically a circle of ~60-100 μm 2 )

However, as shown in Andrasfalvy et al. [43] as long as the scanning time is shorter than the channel deactivation time, stimulation of a larger number of spots can lead to a larger total current than a longer residence time on a single spot. In turn when a certain amount of current is desired (e.g. to fire a cell) this enhancement in efficiency can be used to lower the laser power, reduce the out-offocus light and improve axial resolution. www.sciencedirect.com

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For experiments requiring the delivery of sequences of varying spatiotemporal excitation patterns, that is, n > 1, the refreshing rate, Re, of the LC-SLM has to be taken into account. In that regard in parallel light sculpting the temporal resolution for sequential patterning – that is the time needed to generate a sequence of n patterns – is given by Ts = n  Rt + (n  1)  1/Re ffi n  (Rt + 1/Re) and it is independent of the size of the excitation pattern. For scanning approaches no refreshing time is needed between successive positions and Ts = n  N  (Rt + St). This value can be comparable or less than what is achievable with parallel approaches if: N  (Rt + St)  (Rt + 1/Re). As discussed in [51], more generally a main difference between the serial and parallel approaches is related to potential biological damage, how light is delivered and the mechanism of non-linear excitation. While serial excitation requires moderate average powers but high peak intensity parallel excitation schemes demand higher average powers but modest peak intensities. Higher peak intensities primary leads to non-linear photodamage, while higher average powers result in heating in the brain. Based on the specific dominating mechanisms of the photo-induced damage in a given brain tissue again and the experimental settings (e.g. in vivo stimulation or stimulation in acute brain slice) the parallel or the serial scheme might prove advantageous.

Conclusions and outlook A number of the limitations inherent to both the serial and the parallel excitation scheme, including the required higher power levels for parallel stimulation of large neuronal populations could be overcome by improvements of the optogenetic actuators. Here the development of new constructs with a higher conductance, as is currently underway in different laboratories would have a synergetic impact on the development of optical tools for patterned light stimulation. Also the use of amplified laser systems will permit increasing the excitation field and penetration depth for patterned light stimulation as recently shown for widefield TF imaging [52]. Finally, there is evidence that the axial resolution and shapes of excitation patterns generated by combining TF with low-NA Gaussian beams [53], holographic beams or beams generated with the GPC method [54] are extremely resistant to scattering which hold great promise for application of stimulation patterns at larger depths and in vivo. The light manipulation techniques discussed here and others that can be expected to become available in near future are adding a new dimension to all optimization efforts on the labels’ end by exploiting the optical knob in optogenetics. These efforts represent an important step towards an all optical neuronal control and read out. www.sciencedirect.com

Acknowledgements We would like to thank Simon Rumpel, Andrew Straw and David Tank for their comments on the manuscript. VE is supported by the Human Frontier Science Program (RGP0013/2010) and the ‘‘Fondation pour la Recherche Me´dicale’’ (FRM e´quipe). AV is supported by a grant from Wiener Wissenschafts-, Forschungs- und Technologiefonds (WWTF).

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2. Scanziani M, Hausser M: Electrophysiology in the age of light.  Nature 2009, 461:930-939. A review illustrating the importance of optical methods in neuroscience. 3. Miesenbock G: The optogenetic catechism. Science 2009,  326:395-399. One of the first complete review on optogenetic actuators and sensors. 4. Deisseroth K: Optogenetics. Nat Methods 2011, 8:26-29.  A recent review of the optogenetic techniques and future developments. 5.

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