Science Journals — AAAS - eaao5520.full.pdf

OPTICS

The Authors, some

rights reserved;

exclusive licensee

American Association

for the Advancement

of Science. No claim to

original U.S. Government

Works. Distributed

under a Creative

Commons Attribution

NonCommercial

License 4.0 (CC BY-NC).

Deep tissue optical focusing and optogenetic

modulation with time-reversed ultrasonically

encoded light

Haowen Ruan,

* Joshua Brake,

* J. Elliott Robinson,

Yan Liu,

Mooseok Jang,

†

Cheng Xiao,

‡

Chunyi Zhou,

‡

Viviana Gradinaru,

Changhuei Yang

1§

Noninvasive light focusing deep inside living biological tissue has long been a goal in biomedical optics. However,

the optical scattering of biological ti

ssue prevents conventional optical systems from tightly focusing visible light

beyond several hundred micrometers. The recently developed wavefront shaping technique time-reversed ultra-

sonically encoded (TRUE) focusing enables noninvasive light delivery to targeted locations beyond the optical dif-

fusion limit. However, until now, TRUE focusing has onl

y been demonstrated inside nonliving tissue samples. We

present the first example of TRUE focusing in 2-mm-thick livi

ng brain tissue and demonstrate its application for opto-

genetic modulation of neural activity in 800-

m

m-thick acute mouse brain slices at a wavelength of 532 nm. We found

that TRUE focusing enabled precise control of neuron firing a

nd increased the spatial resolution of neuronal excitation

fourfold when compared to conventional lens focusing. Th

is work is an important step in the application of TRUE

focusing for practical biomedical uses.

INTRODUCTION

Optical methods are widely used across biomedical research, as well as

for the diagnosis and treatment of disease, yet the ability to monitor

and modulate biological processes at depth is conventionally limited

by light scattering caused by the heterogeneous optical properties of

biological samples. For example

, a 532-nm photon experiences an av-

erage of nearly 40 scattering events as it travels through 1 mm of mouse

brain tissue (scattering mean free path

≈

m) (

), which exem-

plifies why the formation of an optical focus in typical tissue samples is

often limited to depths of a few hundred micrometers. To focus light

deeper inside the tissue, wavefront shaping or wavefront engineering

methods (

–

) have been developed that counteract the effects of

optical scattering by modulating the incident light field so that the scat-

tered light controllably interferes at locations of interest to form tight

foci.Thisclassofmethodsprovidesanadvantageovertechniquesthat

discard scattered light as noise, such as confocal microscopy, because

the probability of photons being unsc

attered (that is, ballistic in na-

ture) decays exponentially with increasing depth. The incorporation

of scattered photons enables light focusing beyond the optical diffu-

sion limit where the propagation directions of the photons become

random (

). Furthermore, because wavefront shaping techniques ac-

tively control scattered light, they offer direct optical modulation, an ad-

vantage over other optical imaging techniques such as photoacoustic

tomography (

) and diffuse optical tomography (

), which enable

deep tissue imaging, but cannot focus

light to a particular location for

improved light delivery.

The ability to manipulate scatter

ed photons to create a light focus

at depth with wavefront shaping is due to the elastic, deterministic

nature of optical scattering, which scrambles but does not eliminate

the information contained within a light field (

). Thus, if one could

discern the positions and scattering profile of the scatterers within the

medium, it would be possible to tailor an incident wavefront to opti-

mally couple light to any point in the tissue. This process can be sim-

plified by mapping the optical phase and/or amplitude relationship

between the input plane outside the sample and the targeted plane

inside, which can be accomplished through feedback-based approaches

(

), transmission matrix measurement (

–

), or op-

tical time reversal (optical phase conjugation) (

–

). Among

these, optical phase conjugation is well suited for optical focusing in

living tissue applications because it allows for measurement of the

phase relationship between the target focus and the wavefront solu-

tion on the input plane in parallel, th

us producing the fastest focusing

speeds (

–

). This feature helps to overcome challenges posed by

living tissue dynamics, which requi

re that the wavefront shaping sys-

tem obtain and playback the wavefront solution before the scatterers

’

configuration in the tissue changes (

–

). When used with a guide-

star (

), a method for tagging photons that traverse a desired location

within the biological sample, the o

ptical phase conjugation approach

can create a phase conjugate wavefront that forms a focus at the guide-

star

location.

Several guidestar mechanisms have been developed that enable the

generation of appropriate input wave

fronts. These include fluorescent

(

), nonlinear optical (

–

), kinetic (

), photoacoustic

(

–

), ultrasonic (

–

), magnetic (

), and microbubble (

)en-

coded mechanisms. Of these, ultras

ound offers the advantage of being

noninvasive, freely addressable within the volume of interest, and com-

patible with optical phase conjugation because it generates coherent

tagged light. Time-reversed ultrasonically encoded (TRUE) focusing is

a wavefront shaping technique that combines optical phase conju-

gation with the ultrasound guidestar to enable light focusing at depths

beyond the optical diffusion limit

with ultrasonic resolution (~30

(

–

The application of TRUE focusing t

o living systems would be ben-

eficial to many fields of study, inclu

ding neurobiological research, in

which visible light is routinely used for both monitoring activity with

Department of Electrical Engineering, California Institute of Technology, 1200 East

California Boulevard, Pasadena, CA 91125, USA.

Division of Biology and Biological

Engineering, California Institute of Technology, Pasadena, CA 91125, USA.

*These authors contributed equally to this work.

†

Present address: Department of Physics, Korea University, 145 Anam-ro, Seongbuk-gu,

Seoul 02841, South Korea.

‡

Present address: School of Anesthesiology, Xuzhou Medical University, Xuzhou,

Jiangsu 221004, China.

§Corresponding author. Email: chyang@caltech.edu

SCIENCE ADVANCES

RESEARCH ARTICLE

Ruan

et al

Sci. Adv.

2017;

:eaao5520 8 December 2017

1of9

on December 12, 2017

http://advances.sciencemag.org/

Downloaded from

genetically encoded neural activity indicators (

)andcontrolling

activity via optogenetic actuators (

). Although neurophotonic tech-

niques that use multiphoton excitation (

–

) and adaptive optics

(

) have extended the depths of optical access in vivo, focusing

light noninvasively in the multiple scattering regime in living brain tis-

sue remains largely unexplored. Because of the strongly scattering na-

ture of brain tissue, light delivery during optogenetic manipulation still

requires the use of invasive, implanted optical fibers to reach targets in

deep brain regions (

). Because TRUE focusing allows for an optical focus

to be formed noninvasively with the ability to freely move the focus within

the tissue to target different regions of interest, it is particularly well

suited for optogenetic

modulation. Here, we describe the design and

application of an integrated TRUE focusing and patch clamp electro-

physiology system for simultaneous o

ptogenetic stimulation and neural

activity monitoring within living brain tissue ex vivo. We first demon-

strate light focusing through up to 2-

mm-thick living brain tissue using

diffuse photons with a wavelength of 532 nm. Then, by performing

patch clamp recordings in 800-

m-thick acute brain slices and using

optogenetically evoked photocurrents as a readout, we demonstrate that

TRUE focusing increases the spatial r

esolution of neuronal excitation

by four times compared to that of conventional focusing at a wave-

length of 532 nm. This result represen

ts the first demonstration of TRUE

focusing in living brain tissue and is an important step in the translation

of wavefront shaping methods into practical optical tools for in vivo

applications, including optogenetics.

RESULTS

System design and operating principles

To achieve TRUE focusing in living brain tissue ex vivo, we designed

and implemented a digital optical p

hase conjugation (DOPC) system

(

) for TRUE focusing (figs. S1 and S2) that included an integrated

patch clamp electrophysiology head stage and amplifier for neuro-

physiological measurements, as well as a removable differential in-

terference contrast (DIC) microscope for neuron visualization while

whole-cell recordings were being o

btained (Fig. 1). A customized sam-

ple chamber was designed that allow

ed acute brain slices to rest hori-

zontally while constantly perfused with carbogenated artificial cerebral

spinal fluid (aCSF). Because this set

up limited the orientations of the

TRUE focusing light path and the ultrasound transducer to oblique angles,

we illuminated the slice at a 45° angle with the ultrasound transducer

positioned orthogonal to the TRUE light beam to maximize the mod-

ulation efficiency. To allow for the use of high numerical aperture (NA)

lenses (for the DIC microscope objective, collection lens, and ultra-

sound transducer) with relatively short working distances to be op-

erated within the limited available space, the observation objective,

collection lens with lens tube, and ultrasound transducer were placed

on computer-controlled motorized stages so they could be precisely

translated in and out of the bath. To prevent fluctuations of the per-

fusion fluid surface from influencing the wavefront measurement, we

attached a glass window to the lens tube and immersed it in the aCSF

solution. This normally incident design also avoids unnecessary refrac-

tion at the aCSF-air interface. Similarly, the bottom of the chamber

was also designed with a 45° chamber-air interface, which minimizes

the effects of refraction and helps with optical alignment.

The creation of a TRUE focus involved sequential wavefront record-

ing (Fig. 2A1) and playback (Fig. 2A2) steps. In the recording step, a

high-frequency (50 MHz) ultrasound field was focused to the location

of interest while a probe light beam generated by a pulsed 532-nm

laser illuminated the sample. Because of the acousto-optic effect,

the frequency of a portion of the ligh

t passing through the ultrasound

focus was shifted by the ultrasound frequency. The field of the scat-

tered, ultrasound-tagged light was measured by the camera of the

DOPC system using interferometry (

). Then, in the playback step,

thephaseconjugateversionofthephasemapoftheultrasound-tagged

light was displayed on the spatial light modulator (SLM) of the DOPC

system and used to create the playba

ck light field. Following the prin-

ciple of time reversal, this playback

beam scattered in a time-reversed

fashion and formed an optical focus

at the location of the ultrasound

focus. Our TRUE focusing system d

escribed here relied on a digital

wavefront recording and playback engine (

), which, compared

with analog TRUE fo

cusing systems (

), allowed for measured wave-

fronts to be played back at a light intensity far greater than that of the

measured wave (

The average intensity of the TRUE focus compared to the back-

ground intensity for phase-only modulation of the wavefront is given

by Eq. 1 (

)

phase

‐

only

where

is the number of optical modes controlled by the SLM and

is the number of optical modes (speckle grains) within the ultrasound

Fig. 1. Custom TRUE focusing and electrophysiological recording system.

The

custom TRUE focusing system combined a DO

PC system with a patch clamp electro-

physiology amplifier and headstage. Acute brain slices were held in a custom perfusion

chamber that contained warmed, carboge

nated aCSF. The TRUE light beam illumi-

nated the tissue at an oblique 45° angle, and

the borosilicate patch pipette electrode

was used for neurophysiological measurements. (

) A DIC microscope was included for

neuron visualization during patch clamping. (

) The TRUE focusing system allowed

light to be sharply focused through the brain slice. (

)Aclose-upimageoftheTRUE

focus on a patched neuron. Scale bars, 20

min(A)and50

min(B).

SCIENCE ADVANCES

RESEARCH ARTICLE

Ruan

et al

Sci. Adv.

2017;

:eaao5520 8 December 2017

2of9

on December 12, 2017

http://advances.sciencemag.org/

Downloaded from

focus. The size of the TRUE focus along the ultrasound beam lateral

direction is dictated by the diffraction-limited focused ultrasound beam

diameter, and the size along the ultrasound beam axial direction is

determined by the ultrasound and laser pulse widths. To enhance the

spatial resolution and contrast of TRUE focusing, we used an iterative

TRUE focusing scheme (

–

), where the intensity and resolution of

the TRUE focus were iteratively enhanced by repeating the TRUE

focusing procedure using a previously established TRUE focus. A random

phase pattern was displayed on the SLM to initiate the iterative TRUE

focusing process. Rather than using two DOPC systems as previously

demonstrated (

), we designed and implemented the iterative TRUE

focusing system in transmission mo

de using a single DOPC system (see

MaterialsandMethodsandfigs.S1andS2fordetaileddescriptions).

A comparison between TRUE and conventional focusing

in living brain slices

To test the performance of our system, we prepared acute brain slices

(300 to 2000

m) that contained the medial

prefrontal cortex (mPFC)

from C57Bl/6J mice using a vibrating microtome as previously de-

scribed (

). Then, we placed the slices in our optical setup and

recorded the light intensity profile through the slices formed by our

TRUE focusing system (Fig. 2, A1 and A2) and a conventional lens (Fig.

2B). As predicted, the conventional f

ocusing lens failed to form a tight

optical focus and demonstrated a light profile that broadened as the

brain slice thickness was increased due to the strong scattering nature

of the tissue (Fig. 2C, top row). While

a visible envelope of the intensity

profile was observed when light was

conventionally focused through a

500-

m-thick slice, the lateral width of the focus profile was signif-

icantly increased from the diffraction-limited focus size of ~1

m(the

NA of the focusing lens was 0.25). The size of the conventional focus

continued to broaden as slice thick

ness was increased, and no discern-

ible focus envelope was visible within the 580 × 580

–

field of view

in the 1000-

m or thicker slices.

In contrast, TRUE focusing was able

to maintain a lateral resolution

defined by the size of the ultrasoun

d focus, decoupling the size of the

focus from the focusing depth (Fig.

2C, bottom row). Our system used

a high-frequency ultrasound tra

nsducer with a 50-MHz nominal cen-

ter frequency, a 6.35-mm aperture

, and a 12.7-mm focal length. The

theoretical beam diameter (

−

6 dB) for this configuration was ~80

and the calibrated waveform duration (

−

6 dB) was 37.4 ns, correspond-

ing to a pulse length of 55.3

m. The region of ultrasound-modulated

light along the axial direction of the ultrasound beam was also deter-

mined by the combination of the ultrasound pulse length and the laser

pulse duration, which is 7 ns. Using the iterative TRUE focusing

method enabled the TRUE focus to be tightened (

) to achieve a

focus with an average FWHM spot size of 27.4

m across tissue thick-

nesses from 500 to 2000

m (see method S1 for calculation). In con-

trast, the FWHM of the conventional focus broadened from ~350

mat

a slice thickness of 500

m to approximately 2100

matathicknessof

2000

m (Fig. 2D). It should be noted that the effective thicknesses in

the TRUE focusing case are larger than the physical thicknesses of the

slices due to the 45° incident angle of the TRUE focusing beam. These

results demonstrate the ability of

TRUE focusing to overcome optical

scattering to create high-resolution optical foci in living brain slices up

to 2000

m thick, which, unlike those formed by conventional focusing,

do not significantly broaden with increased sample thickness.

0.5

1.5

Focus FWHM versus thickness

Tissue thickness (mm)

Focus FWHM (

μ

Conventional

TRUE

Camera

SLM

Beam splitter

Camer

SLM

Conventional

TRUE

Conventional focusing

TRUE focusing

2000

μm

Lens

Plane wave

Objective

Brain slice

DOPC

Light beam

Ultrasound beam

DOPC

Recording

Playback

1500

μm

1000

μm

800

μm

500

μm

Fig. 2. A comparison of TRUE focusing and conventional focusing.

(

) The recording (A1) and playback (A2) procedures used to focus light through the slice onto

its top surface with TRUE focusing. (

) Diagram of the experimental setup for measuring the light intensity distribution of the focus on the top surface of the brain slice

achieved using a conventional lens illuminating the brain slice from below. A tube lens and a camera used together with the objective are not shown. (

) Images of the

conventional and TRUE focus profile through living brain tissue slices (500, 800, 1000, 1500, and 2000

m thick). (

) Full width at half maximum (FWHM) focal spot sizes

for the conventional and TRUE foci as a function of tissue thickness. Error bars represent the SD of five measurements taken at different locations. Scale bar, 100

SCIENCE ADVANCES

RESEARCH ARTICLE

Ruan

et al

Sci. Adv.

2017;

:eaao5520 8 December 2017

3of9

on December 12, 2017

http://advances.sciencemag.org/

Downloaded from

Application of TRUE focusing for optogenetic manipulations

AfterdemonstratingtheabilityofTRUEfocusingtoovercomeoptical

scattering and produce light foci in thick acute brain slices, we next

soughttodemonstratetheadvantageofTRUEforoptogeneticmanipu-

lation compared with conventional focusing using a neurophysiological

readout. Optogenetics, in which engineered light-gated ion channels or

pumps are used to manipulate cellular activity with high spatial and

temporal precision using visible lig

ht, has become relatively ubiquitous

in basic neurobiological research due to its ability to convert differences

in light intensity into graded electrophysiological signals (

). Although

awiderangeofoptogeneticactuatorsareavailableforneuralexcitationor

inhibition with diverse excitation spectra spanning the visible spectrum,

we used the excitatory, red-shifted opsin bReaChES for our experiments

because its excitation peak was well matched with our laser source (532

nm) (

). To prepare samples for testing, we performed stereotaxic in-

jections of an adeno-associated viral vector carrying the bReaChES

transgene (AAV-DJ-CaMKII-bReaChES-TS-YFP) into the mPFC of

C57Bl/6J mice (Fig. 3A). After waiting 4 weeks for surgical recovery

and transgene expression, we prepared acute brain slices for simulta-

neous electrophysiological recording and optical testing. Animal hus-

bandry and all experimental procedures involving animal subjects

were approved by the Institutional Animal Care and Use Committee

(IACUC) and by the Office of Laboratory Animal Resources at the

California Institute of Technology under IACUC protocol 1650. We

characterized the performance of bReaChES in cortical slices by mea-

suring the photocurrent response to a wide range of 532-nm light

intensities delivered through the DIC objective in voltage clamp

mode. Similar to its parent opsin ReaChR (

), bReaChES displayed

a nonlinear increase in photocurre

nt response that saturated at an in-

tensity of approximately 10 mW/mm

(Fig. 3B). During these experi-

ments, the average maximum photocurrent across the 10 cells studied

was 1047 pA.

To demonstrate the capability of TRUE focusing for neural mod-

ulation, whole-cell patch clamp recordings were obtained from layer

II/III neurons just below the superf

icial surface of mPFC slices using

borosilicate glass patch pipette electrodes visualized under DIC mi-

croscopy (Fig. 1 and fig. S1C). Although TRUE focusing through slices

up to 2 mm thick was achieved, maxi

mum slice thickness during our

optogenetic experiments was limited to 800

m, because neurons in

thicker slices were difficult to visualize with DIC microscopy and were

generally less healthy, which negatively affects recording and data

quality. Because target neurons were located close to the surface of

the brain slice for visualization, the DOPC playback beam illuminated

from the bottom of the slice traversed almost the entire sample thick-

ness, which is much larger than the optical diffusion limit of the acute

mouse brain slice (~200

mat532nm)(

). Moreover, because the

incident angle was 45° (Fig. 3C), the effective thickness for TRUE fo-

cusing was even larger than the phys

ical slice thickness. Once a whole-

cell recording was successfully init

iated, the DIC microscope objective

was removed, and the lens tube and ultrasound transducer were

lowered into the bath. To ensure colocalization of the ultrasound focus

with the pipette tip, we used pulse-echo ultrasound to form an image of

the glass pipette tip (fig. S3) and moved the ultrasound transducer to

focusontheendofthetipwhereatargetneuronwaslocated.This

approach allowed precise targeting of the TRUE focus to the recording

neuron to maximize light delivery during optogenetic stimulation.

Next, we measured the photocurrent response that was elicited by

the TRUE focus; as a control, we created a

“

no wavefront shaping

”

con-

dition by shifting the wavefront solution on the SLM by 100 pixels in

each lateral direction, which generally approximated the laser back-

ground intensity. In this case, the TRUE focus outperformed the no

shaping condition, evoking a larger photocurrent due to enhanced

light intensity at the focus (Fig. 3D, left). The photocurrent enhance-

ment factor, defined as the ratio between the difference of the photo-

current with and without TRUE focusing and the photocurrent without

TRUE focusing, was on average 30% (

= 6) in 800-

m-thick brain

slices, which was similar in magn

itude to the enhancement observed

in 300- and 500-

m slices (fig. S4). To verify the effect of the ultrasound

guidestar, we turned off the ultrasound and repeated the same

procedure. In this case, because there was no guidestar for the system

to focus to, no TRUE focus was form

ed, resulting in a smaller evoked

photocurrent and no observed firing events (fig. S5). Because the

presence of the ultrasound field could potentially alter neural activity,

we verified in several neurons that focused ultrasound alone in the

absence of light failed to evoke any observable current in voltage

clamp or alter neuronal excitability in current clamp mode.

We next sought to evaluate the performance of our system by com-

paring the experimentally observed enhancement factor with the ex-

pected enhancement predicted by the technical specifications of our

system and the observed TRUE focus size. The SLM used in the DOPC

system had 2 × 10

pixels, which allows us to focus light through a

highly scattering medium to a singl

e optical mode with an experimen-

tal peak focus intensity to background ratio

of ~1 × 10

.Thisexper-

imental performance means the DOPC system could effectively control

~1×10

optical modes. On the basis of this performance, we were

able to estimate the intensity enhancement at the ultrasound focus

using Eq. 1. Because our system produced a TRUE focus with a FWHM

Action potentials

Electrode

Neurons

Camera

SLM

DOPC

Beam splitter

800 μm

TRUE No shaping

TRUE

No shaping

100 pA

1 s

Membrane voltage

Photocurrent

20 mV

1 s

AAV-DJ-CaMKII-bReaChES-TS-YFP

mPFC

532 nm

Light intensity (mW/mm

)

0.2

0.4

0.6

1.0

0.8

Normalized photocurren

bReaChES photocurren

n

= 10

Fig. 3. Experiment design, opsin characterization, and demonstration of photo-

current and firing modulation via TRUE focusing.

(

) An AAV vector was used to

stereotaxically deliver the bReaChES transgene to the mPFC. (

) Characterization

of normalized photocurrent response versus light intensity. The average maxi-

mum photocurrent across the 10 cells studied was 1047 pA. (

) Diagram illustrat-

ing the experimental scheme used to demonstrate the ability of TRUE focusing to

elicit action potentials through 800-

m-thick living mouse brain tissue. (

) Repre-

sentative traces demonstrating elicited photocurrent and membrane voltage changes

achieved with and without TRUE focusing.

SCIENCE ADVANCES

RESEARCH ARTICLE

Ruan

et al

Sci. Adv.

2017;

:eaao5520 8 December 2017

4of9

on December 12, 2017

http://advances.sciencemag.org/

Downloaded from

diameter of ~27

m, the number of modes

inside the focus was

~1 × 10

, which corresponded to a predicted intensity enhancement

factor at the ultrasound focus of a

pproximately 2. Because the photo-

current enhancement was not proportional to light intensity (Fig. 3B),

we predicted that the photocurrent

enhancement factor would be less

than 1, which was consistent with ou

r data. Despite the observed en-

hancement, the laser power could be adjusted so that the TRUE focus

elicited time-locked cell firing, whereas the no shaping condition could

not elicit action potentials (Fig. 3D, right).

Improved spatial resolution of optogenetic stimulation

using TRUE focusing

After demonstrating that TRUE-focused light could optogenetically

stimulate neurons at depths beyond th

e optical diffusion limit, we com-

pared the spatial resolution of TRUE f

ocusing with that of conventional

lens focusing for optogenetic modulation. The ability for TRUE focusing

to noninvasively enhance the light intensity in a spatially restricted man-

ner is an important benefit compared to other conventional methods for

delivering light into the brain, such as

opticalfibersorlight-emittingdi-

ode (LED) implants, which do not allow for the targeted volume to be

freely moved within the brain after implantation. To quantify the spatial

resolution of TRUE focusing and conventional focusing, we raster-

scanned the focus of each case laterally around a patch-clamped neuron

and recorded the photocurrent magnitude at each scanning position

(Fig. 4, A and B). In both cases, we scanned over a square grid of 9 × 9

points with a 50-

m step size in each dimension on the horizontal

plane. For conventional focusing,

the 780-nm wavelength DIC illumi-

nation LED was replaced with the 532-nm wavelength laser source

delivered via a single-mode optical fiber whose tip was imaged to the

plane of the targeted neuron to form a focus (fig. S1D). The position

of the focus was calibrated using the observation microscope before

placing the brain slice in the chamber, and the focus was raster-scanned

on the horizontal plane during whole-cell recordings (fig. S1, C and D).

The normalized photocurrent enhancement was calculated at each

scanning position and used to construct interpolated two-dimensional

(2D) scan maps (Fig. 4, C and D). Fitting the conventional lens scan

map with a 2D Gaussian function yielded respective FWHMs of 393

and 536

minthe

and

dimensions. In contrast, the FWHMs for the

TRUE focusing scan were 99 and 71

minthe

and

dimensions.

Because of the scattering and diffusion of the conventional illumina-

tion, the spatial extent of the evoked photocurrent enhancement with

conventionalilluminationwasnearlyfourtimesbroaderthanthatob-

tained with TRUE focusing, thus confirming the utility of TRUE

focusing for precise spatial focusing at depth beyond the optical diffu-

sion limit.

DISCUSSION

Overcoming optical scattering to noninvasively extend the depth at

which light can be tightly focused inside living biological samples in

clinical and research settings is of great interest to practitioners and re-

searchers alike. Here, we developed a TRUE focusing system that allowed

us to focus light at depth in ex vivo brain tissue with a spatial resolution

that significantly outperformed conventional lens focusing. By integrat-

ing a patch clamp electrophysiology headstage and amplifier into the

TRUE focusing system, we were able to monitor neural activity during

optogenetic stimulation with the TRUE focus. Using neurophysiological

signals as a readout, we confirmed that TRUE focusing can be used to

control neural activity in thick tissue samples in a spatially restricted

manner. Because optogenetic manipul

ations currently require the surgi-

cal implantation of invasive optical fibers for light delivery below the

most superficial brain regions (

), we believe that our findings using

TRUE focusing will inform future efforts to develop this technology

for noninvasive optogenetic stimulation and/or fluorescent imaging in

vivo with the spatial resolution requir

ed for precise targeting of individ-

ual neurons or neuron ensembles.

Multiphoton microscopy is capable of obtaining clear images at

depths of 800

m and is promising for neuromodulation at that depth.

However, the fundamental working depth of this technique is limited by

the number of unscattered or weakly scattered photons, which de-

creases exponentially with depth. In contrast, the TRUE focusing tech-

nique is able to focus light beyond the ballistic photon regime. The

addressable depth of the TRUE focus

ing technique demonstrated in this

set of optogenetic experiments was

limited by the penetration depth of

the DIC microscope illumination ne

cessary to visualize neurons during

the initiation of patch clamp record

ings, as well as the viability of the

neurons in thick tissue. Although fluorescent activity indicators, such as

the GCaMP family of proteins (

), would provide a viable activity

readout in thicker tissue samples, these tools were not practical for

use here given that the excitation wavelength for calcium indicators is

likely to simultaneously excite neurons with opsins that match the

operating wavelength of the TRUE focusing system (532 nm). In the

future, this problem could be solved by decoupling the wavelength

for TRUE focusing and optogenetic excitation from that for calcium

indicator excitation. It would also be valuable to explore the maximum

penetration depth of TRUE focusing for optogenetics, even if it would

require minimally invasive methods in vivo such as optical fiber in-

sertion for signal readout.

800 μm

Electrode

Lens

Brain slice

Photocurrent

Conventional

TRUE

0.2

0.4

0.6

0.8

Normalized

photocurrent enhancement

Camera

SLM

Beam splitter

DOPC

Ultrasound

scanning

Fig. 4. Spatial resolution of optogenetic stimulation achieved by conventional

versus TRUE focusing.

Experimental configuration for photocurrent scan map

generation using conventional focusing (

)andTRUEfocusing(

). The normalized

photocurrent enhancement as a function of lateral focal scanning position for

conventional (

) and TRUE focusing (

). Scale bar, 100

SCIENCE ADVANCES

RESEARCH ARTICLE

Ruan

et al

Sci. Adv.

2017;

:eaao5520 8 December 2017

5of9

on December 12, 2017

http://advances.sciencemag.org/

Downloaded from

Unlike in ex vivo tissue preparations

where cell viability is a limiting

factor for tissue thickness, the focusing depth during in vivo applica-

tions is limited by the guidestar efficiency. As we focus deeper into

tissue, fewer photons from the guidestar can be measured, not only be-

cause the detected portion of light fr

om the guidestar is reduced but also

because of a decrease in the modulation efficiency due to ultrasound

attenuation. Although the DOPC system works even when the mea-

sured phase map has less than a photon per degree of freedom (that

is, SLM or camera pixel) (

), the presence of shot noise due to the

much higher unmodulated light intensity will fundamentally limit

the penetration depth (

). Additional guidestar aids, such as micro-

bubbles, can help improve the tagging efficiency significantly (

)but

sacrifice the freely addressable and n

oninvasive nature of the ultrasound

guidestar. In the future, it will be important to optimize the intensity

of the measured ultrasound-modulated light to extend the penetra-

tion depth.

Another goal for future developments of TRUE focusing for opto-

genetic simulation is improved photocurrent enhancement. Using

whole-cell recordings, we observed a photocurrent enhancement of ap-

proximately 30% compared to the no shaping condition, which was

consistent with predicted values but will require improvement before

TRUE focusing is feasible for widespread use in optogenetic applica-

tions. The avenues to improve the focusing contrast are based on the

variables in Eq. 1. From this equation, we can see that to enhance the

focusing contrast, we can either increase

, the number of controllable

modes, or decrease

, the number of optical modes within the ultra-

sound focus. One way to reduce

is by reducing the size of the ul-

trasound focus by increasing the operating frequency and the NA of the

ultrasound transducer. However, hig

h-frequency ultrasound has a very

limited penetration depth. Furthermore, because the goal is to enhance

the light intensity delivered to the neuron soma, shrinking the size of the

focus beyond the size of the cell will not necessarily lead to further im-

provements in photocurrent enhanc

ement, although this strategy may

allow for finer resolution targeting of neuronal subcompartments, such

as individual dendrites or synaptic in

puts.Anotherstrategyistoincrease

the size of the optical modes by shifting to longer wavelengths, although

opsins sensitive to infrared or near-infrared wavelengths will need to

be further refined before they are practical for single photon in vivo ap-

plications (

). A more feasible avenue to improve the TRUE focus

contrast is to increase the number of controllable optical modes,

.This

can be achieved by scaling up the number of SLM pixels, which will also

benefit other general applications across the wavefront shaping field.

For example, increasing the number of SLM pixels

by 10 times will

result in a focus intensity

–

background ratio

≈

12, which is suf-

ficient for many practical applications.

To translate wavefront shaping into practical tools for in vivo ap-

plications, we also need to address the challenge of the optical decor-

relation of living tissue. The dynamic nature of living tissue causes

decorrelation of the optical wavefronts, so to effectively focus light

inside living tissue, the system response time must be shorter than

the decorrelation time of the tissue. For acute brain slices less than

2 mm in thickness, this decorrelation time is on the order of several

seconds (

),whichislongerthanthecurrentTRUEfocusingspeed

(0.6 s; see fig. S2). However, the de

correlation time drops to the order

of 1 ms for in vivo applications due to blood flow, cardiac motion,

breathing, etc. (

). To increase the response speed of wavefront

shaping systems, digital micromirror devices and ferroelectric liquid

crystal

–

based SLMs have been used to achieve high-speed DOPC with-

in 10 ms (

), which is ultimately limited by the need to read out and

transfer data to a computing device such as a personal computer or

an embedded system to compute the appropriate wavefront solution.

We expect that solving these problems will require an integrated

wavefront

shaping system that combines the wavefront sensing and

modulation devices into a single device (

). This design will allow

for control over an increased number of optical modes in a scalable

way without sacrificing the operation speed, because wavefront cal-

culations can be performed in parallel on a per pixel basis, minimiz-

ing data transfer and computation time. The development of such an

integrated wavefront sensing and modulation platform will increase

the achievable enhancement factors. Simultaneously, it will remove

many of the challenges that limit the widespread adoption of wavefront

shaping techniques, such as the difficulty of designing and aligning the

complex optical system (

), opening the door for more scientists to

incorporate wavefront shaping into their optical technologies for bio-

medicine and beyond.

MATERIALS AND METHODS

Acute brain slice preparation

Stereotaxic injection of AAV-DJ-CaMKII-bReaChES-TS-YFP was

used to deliver the opsin transgene into the mPFC in adult mice and

allowed to express for 3 to 4 weeks before the experiments were con-

ducted. On the day of each experiment, acute brain slices (300 to

2000

m) that contained the infralimbic and prelimbic cortices were

prepared with a vibrating microtome after euthanasia and transcar-

dial perfusion with ice-cold cutting solution, as previously described

(

). Slices were recovered in 32°C, carbogenated aCSF for 1 hour

before the start of each recording. Recordings were performed using a

potassium gluconate internal solution in the presence of carbogenated

aCSF that contained 3 mM kynurenic acid to block excitatory post-

synaptic currents. Methods S2 and S3 describe sample preparation

and recording conditions in greater detail.

TRUE focusing system design and integration of the patch

clamp amplifier and head stage

The DOPC system consisted of three major modules (fig. S1, A and

B): a light beam preparation module, a DOPC module, and the patch

clamp electrophysiology amplifie

r/head stage. The light beam prepa-

ration module prepared three light beams for the DOPC system: a

planar reference beam for wavefront recording (R), a sample or playback

beam (S/PB) that illuminated the sample, and a quality assurance

beam (QA) for daily system alignmen

t. All three beams were spatially

filtered, path-length

–

matched, and aligned to the horizontal polariza-

tion direction. For the S/PB beam, w

e used two laser sources, a nano-

second pulsed wave (PW) laser (532-nm wavelength, 7-ns pulse width,

40-kHz repetition rate, and 7-mm cohe

rence length; Navigator, Spectra-

Physics) for TRUE focusing and a con

tinuous wave (CW) laser (532-nm

wavelength; Millennia eV, Spectra-Physics) for optogenetic stimula-

tion, which was modulated by an optical chopper. These two laser

beams were selected by a beam selecting shutter (BSS1) and coupled

to a pinhole-based spatial filter through a beam splitter (BS2). A 4f sys-

tem (L1 and L2) was used to match the beam diameter of the PW laser

to that of the CW laser beam so that they achieved optimum coupling

efficiency through the pinhole. The frequency of the reference beam and

QA beam was modulated by two acousto-optic modulators (AOM;

AFM-502-A1, IntraAction).

The DOPC module used four beam splitters (BS4, BS5, BS8, and

BS9) and two beam selecting shutters (BSS2 and BSS3) to route the

SCIENCE ADVANCES

RESEARCH ARTICLE

Ruan

et al

Sci. Adv.

2017;

:eaao5520 8 December 2017

6of9

on December 12, 2017

http://advances.sciencemag.org/

Downloaded from

S/PB beam into two separated optical loops (loop A and loop B),

which were used sequentially during iterative TRUE focusing operation.

To initialize the iterative proces

s, the SLM (PLUTO, HOLOEYE) dis-

played a random phase map to gener

ate a disordered light field that

mimicked the light inside the sample. The BSS2 and BSS3 were set

to enable DOPC loop A, which directed the light to the sample from

the top surface. The ultrasound-modulated light field was measured

by the camera (Camera 1, pco.edge 5.5, PCO-TECH) of the DOPC sys-

tem, and its conjugated phase map was displayed on the SLM. We

then flipped BSS2 and BSS3 to enable DOPC loop B, which routed

the shaped S/PB beam to the sample in the reversed direction (fig. S1B),

resulting in an initial TRUE focus. The TRUE focus was modulated by

the ultrasound again, and the ultrasound-modulated light was measured

by the camera. Immediately, the SLM was updated and DOPC loop A

was enabled again for the next iteration. By repeating the TRUE focusing

process between these two DOPC loops nine times (nine SLM updates),

we obtained an optimized wavefron

t solution for TRUE focusing. In

this case, DOPC loop B is enabled and we switched the light source

to the CW laser for neural modulation. An amplitude mask (ZB) was

placed on the focal plane of lens L7 to block the zeroth order of the

playback beam, which was not modulated by the SLM. The daily tuning

procedure for the DOPC system can be found in method S4.

The electrophysiological recordi

ng setup had two operating modes,

aneuronpatchingmodeandaneuronstimulationmode(fig.S1,C

and D). In the neuron patching mode (fig. S1C), the collection lens

(L10) and its lens tube and the transducer were translated out of the

chamber, allowing the objective

(40×, LUMPlanFL/IR, Olympus) of a

custom-built DIC microscope to be immersed into the solution for

neuron visualization. We used a 780-nm LED (M780D2, Thorlabs)

as the light source (LS1) to maximize the penetration depth. Once a

whole-cell recording was initiated, we switched to the neuron stim-

ulation mode (fig. S1D) by lifting the objective out of the perfusion

chamber and translating the ultrasonic transducer (PI50, Olympus)

and collection lens L10 and its lens tube down to the chamber. In this

mode, we performed iterative TRUE f

ocusing while photocurrents or

transmembrane potentials were measured via the patch pipette elec-

trode (PP). In the case of conventional focusing, we replaced the LED

source with the 532-nm CW laser source delivered by a single-mode

fiber (SF3) whose tip was imaged to the top surface of the sample.

Weraster-scannedthefocusbyscanningthetipofthesinglemodefiber

SF3 on the focal plane of lens L17 while recording the transmembrane

current at each scanning position. The

electrophysiological signals were

recorded by a computer-controlled patch clamp amplifier (EPC10 USB,

HEKA) and filtered at 10 kHz.

Measurement of the phase map of the ultrasonically

tagged light

A detailed signal flow diagram is shown in fig. S2. We used four-step

phase-shifting holography (

) to measure the phase of the ultra-

sonically tagged light and shifted the phase of the reference beam by

stepping the phase of the signal driving the AOM through 0,

/2,

and 3

/2. Four intensity maps (

,and

) corresponding to

each phase of the reference beam were recorded, and the phase map of

the ultrasonically tagged light was calculated as

= Arg[(

−

(

−

)], where Arg[·] computes the principal value of the argument

of a complex number. Because we used laser pulses with a pulse width

smaller than 20 ns as the light source to ensure fine axial resolution of

the TRUE focus, we needed to carefully design the parameters to elim-

inate the unwanted signal formed by the interference between the

reference beam (R) and the unmodulated sample beam (U), as well

as the unwanted signal formed by the interference between the ultra-

sonically tagged light (T) and U, which would otherwise overwhelm

the real signal formed by the interference between R and T. The for-

mulation to design the timing for these three beams has been de-

scribed previously (

), and a detailed signal diagram is illustrated

in fig. S2.

SUPPLEMENTARY MATERIALS

Supplementary material for this article is available at http://advances.sciencemag.org/cgi/

content/full/3/12/eaao5520/DC1

method S1. Calculation of the focal spot size of TRUE and conventional focusing.

method S2. Viral injection surgery.

method S3. Electrophysiological recordings.

method S4. Daily alignment procedure.

fig. S1. Setup.

fig. S2. Electrical signal flow diagram.

fig. S3. Ultrasound pulse-echo image of the tip of the glass pipette electrode.

fig. S4. Electrophysiological photocurrent traces from neurons in 500- and 300-

m-thick acute

brain slices.

fig. S5. Electrophysiological photocurrent and membrane voltage traces comparing ultrasound

on and off conditions.

REFERENCES AND NOTES

1. S. Schott, J. Bertolotti, J.-F. Léger, L. Bourdieu, S. Gigan, Characterization of the angular

memory effect of scattered light in biological tissues.

Opt. Express

, 13505

–

13516

(2015).

2. M. Mesradi, A. Genoux, V. Cuplov, D. Abi-Haidar, S. Jan, I. Buvat, F. Pain, Experimental and

analytical comparative study of optical coefficient of fresh and frozen rat tissues.

J. Biomed. Opt.

, 117010 (2013).

3. I. M. Vellekoop, A. P. Mosk, Focusing coherent light through opaque strongly scattering

media.

Opt. Lett.

, 2309

–

2311 (2007).

4. A. P. Mosk, A. Lagendijk, G. Lerosey, M. Fink, Controlling waves in space and time for

imaging and focusing in complex media.

Nat. Photonics

, 283

–

292 (2012).

5. R. Horstmeyer, H. Ruan, C. Yang, Guidestar-assisted wavefront-shaping methods for

focusing light into biological tissue.

Nat. Photonics

, 563

–

571 (2015).

6. H. Yu, J. Park, K. Lee, J. Yoon, K. Kim, S. Lee, Y. Park, Recent advances in wavefront shaping

techniques for biomedical applications.

Curr. Appl. Phys.

, 632

–

641 (2015).

7. S. Rotter, S. Gigan, Light fields in complex media: Mesoscopic scattering meets wave

control.

Rev. Mod. Phys.

, 015005 (2017).

8. I. M. Vellekoop, Feedback-based wavefront shaping.

Opt. Express

, 12189

–

12206

(2015).

9. S. M. Popoff, G. Lerosey, M. Fink, A. C. Boccara, S. Gigan, Controlling light through optical

disordered media: Transmission matrix approach.

New J. Phys.

, 123021 (2011).

10. M. Kim, W. Choi, Y. Choi, C. Yoon, W. Choi, Transmission matrix of a scattering medium

and its applications in biophotonics.

Opt. Express

, 12648

–

12668 (2015).

11. L. V. Wang, S. Hu, Photoacoustic tomography: In vivo imaging from organelles to organs.

Science

335

, 1458

–

1462 (2012).

12. D. A. Boas, D. H. Brooks, E. L. Miller, C. A. DiMarzio, M. Kilmer, R. J. Gaudette, Q. Zhang,

Imaging the body with diffuse optical tomography.

IEEE Signal Process. Mag.

,57

–

(2001).

13. Z. Yaqoob, D. Psaltis, M. S. Feld, C. Yang, Optical phase conjugation for turbidity

suppression in biological samples.

Nat. Photonics

, 110

–

115 (2008).

14. J. Yoon, M. Lee, K. R. Lee, N. Kim, J. M. Kim, J. Park, H. Yu, C. Choi, W. D. Heo, Y. K. Park,

Optogenetic control of cell signaling pathway through scattering skull using wavefront

shaping.

Sci. Rep.

, 13289 (2015).

15. R. Sarma, A. G. Yamilov, S. Petrenko, Y. Bromberg, H. Cao, Control of energy density inside

a disordered medium by coupling to open or closed channels.

Phys. Rev. Lett.

117

086803 (2016).

16. S. M. Popoff, G. Lerosey, R. Carminati, M. Fink, A. C. Boccara, S. Gigan, Measuring the

transmission matrix in optics: An approach to the study and control of light propagation

in disordered media.

Phys. Rev. Lett.

104

, 100601 (2010).

17. Y. Choi, T. Daniel Yang, C. Fang-Yen, P. Kang, K. Jin Lee, R. R. Dasari, M. S. Feld, W. Choi,

Overcoming

the diffraction limit using multiple light scattering in a highly disordered

medium.

Phys. Rev. Lett.

107

, 023902 (2011).

18. H. Yu, T. R. Hillman, W. Choi, J. O. Lee, M. S. Feld, R. R. Dasari, Y. K. Park, Measuring large

optical transmission matrices of disordered media.

Phys.Rev.Lett.

111

, 153902 (2013).

SCIENCE ADVANCES

RESEARCH ARTICLE

Ruan

et al

Sci. Adv.

2017;

:eaao5520 8 December 2017

7of9

on December 12, 2017

http://advances.sciencemag.org/

Downloaded from

19. D. B. Conkey, A. M. Caravaca-Aguirre, R. Piestun, High-speed scattering medium

characterization with application to focusing light through turbid media.

Opt. Express

1733

–

1740 (2012).

20. E. N. Leith, J. Upatnieks, Holographic imagery through diffusing media.

J. Opt. Soc. Am.

, 523 (1966).

21. M. Cui, C. Yang, Implementation of a digital optical phase conjugation system and its

application to study the robustness of turbidity suppression by phase conjugation.

Opt. Express

, 3444

–

3455 (2010).

22. C.-L. Hsieh, Y. Pu, R. Grange, D. Psaltis, Digital phase conjugation of second harmonic

radiation emitted by nanoparticles in turbid media.

Opt. Express

, 12283

–

12290

(2010).

23. T. R. Hillman, T. Yamauchi, W. Choi, R. R. Dasari, M. S. Feld, Y. K. Park, Z. Yaqoob,

Digital optical phase conjugation for delivering two-dimensional images through turbid

media.

Sci. Rep.

, 1909 (2013).

24. Y. Shen, Y. Liu, C. Ma, L. V. Wang, Focusing light through biological tissue and

tissue-mimicking phantoms up to 9.6 cm in thickness with digital optical phase

conjugation.

J. Biomed. Opt.

, 085001 (2016).

25. D. Wang, E. Haojiang Zhou, J. Brake, H. Ruan, M. Jang, C. Yang, Focusing through dynamic

tissue with millisecond digital optical phase conjugation.

Optica

, 728

–

735 (2015).

26. Y. Liu, P. Lai, C. Ma, X. Xu, A. A. Grabar, L. V. Wang, Optical focusing deep inside dynamic

scattering media with near-infrared time-reversed ultrasonically encoded (TRUE) light.

Nat. Commun.

, 5904 (2015).

27. Y. Liu, C. Ma, Y. Shen, J. Shi, L. V. Wang, Focusing light inside dynamic scattering media

with millisecond digital optical phase conjugation.

Optica

, 280

–

288 (2017).

28. C. Ma, F. Zhou, Y. Liu, L. V. Wang, Single-exposure optical focusing inside scattering media

using binarized time-reversed adapted perturbation.

Optica

, 869

–

876 (2015).

29. Y. Liu, C. Ma, Y. Shen, L. V. Wang, Bit-efficient, sub-millisecond wavefront measurement

using a lock-in camera for time-reversal based optical focusing inside scattering

media.

Opt. Lett.

, 1321

–

1324 (2016).

30. M. Jang, H. Ruan, I. M. Vellekoop, B. Judkewitz, E. Chung, C. Yang, Relation between

speckle decorrelation and optical phase conjugation (OPC)-based turbidity suppression

through dynamic scattering media: A study on in vivo mouse skin.

Biomed. Opt. Express

–

85 (2015).

31. I. M. Vellekoop, M. Cui, C. Yang, Digital optical phase conjugation of fluorescence in

turbid tissue.

Appl. Phys. Lett.

101

, 081108 (2012).

32. I. M. Vellekoop, E. G. van Putten, A. Lagendijk, A. P. Mosk, Demixing light paths inside

disordered metamaterials.

Opt. Express

,67

–

80 (2008).

33. N. Ji, D. E. Milkie, E. Betzig, Adaptive optics via pupil segmentation for high-resolution

imaging in biological tissues.

Nat. Methods

, 141

–

147 (2010).

34. J. Tang, R. N. Germain, M. Cui, Superpenetration optical microscopy by iterative

multiphoton adaptive compensation technique.

Proc. Natl. Acad. Sci. U.S.A.

109

8434

–

8439

(2012).

35. O. Katz, E. Small, Y. Guan, Y. Silberberg, Noninvasive nonlinear focusing and imaging

through strongly scattering turbid layers.

Optica

, 170

–

174 (2014).

36. I. N. Papadopoulos, J.-S. Jouhanneau, J. F. A. Poulet, B. Judkewitz, Scattering

compensation by focus scanning holographic aberration probing (F-SHARP).

Nat. Photonics

,116

–

123 (2017).

37. E. H. Zhou, H. Ruan, C. Yang, B. Judkewitz, Focusing on moving targets through scattering

samples.

Optica

, 227

–

232 (2014).

38. C. Ma, X. Xu, Y. Liu, L. V. Wang, Time-reversed adapted-perturbation (TRAP) optical

focusing onto dynamic objects inside scattering media.

Nat. Photonics

, 931

–

936 (2014).

39. F. Kong, R. H. Silverman, L. Liu, P. V. Chitnis, K. K. Lee, Y. C. Chen, Photoacoustic-guided

convergence of light through optically diffusive media.

Opt. Lett.

, 2053

–

2055 (2011).

40. A. M. Caravaca-Aguirre, D. B. Conkey, J. D. Dove, H. Ju, T. W. Murray, R. Piestun, High

contrast three-dimensional photoacoustic imaging through scattering media by localized

optical fluence enhancement.

Opt. Express

, 26671

–

26676 (2013).

41. T. Chaigne, O. Katz, A. C. Boccara, M. Fink, E. Bossy, S. Gigan, Controlling light in scattering

media non-invasively using the photoacoustic transmission matrix.

Nat. Photonics

–

64 (2014).

42. P. Lai, L. Wang, J. W. Tay, L. V. Wang, Photoacoustically guided wavefront shaping for

enhanced optical focusing in scattering media.

Nat. Photonics

, 126

–

132 (2015).

43. J. W. Tay, P. Lai, Y. Suzuki, L. V. Wang, Ultrasonically encoded wavefront shaping for

focusing into random media.

Sci. Rep.

, 3918 (2014).

44. X. Xu, H. Liu, L. V. Wang, Time-reversed ultrasonically encoded optical focusing into

scattering media.

Nat. Photonics

, 154

–

157 (2011).

45. Y. M. Wang, B. Judkewitz, C. A. DiMarzio, C. Yang, Deep-tissue focal fluorescence imaging

with digitally time-reversed ultrasound-encoded light.

Nat. Commun.

, 928 (2012).

46. K. Si, R. Fiolka, M. Cui, Fluorescence imaging beyond the ballistic regime by ultrasound-

pulse-guided digital phase conjugation.

Nat. Photonics

, 657

–

661 (2012).

47. H. Ruan, T. Haber, Y. Liu, J. Brake, J. Kim, J. M. Berlin, C. Yang, Focusing light inside

scattering media with magnetic-particle-guided wavefront shaping.

Optica

, 1337

–

1343

(2017).

48. H. Ruan, M. Jang, C. Yang, Optical focusing inside scattering media with time-reversed

ultrasound microbubble encoded light.

Nat. Commun.

, 8968 (2015).

49. B. F. Fosque, Y. Sun, H. Dana, C.-T. Yang, T. Ohyama, M. R. Tadross, R. Patel, M. Zlatic,

D. S. Kim, M. B. Ahrens, V. Jayaraman, L. L. Looger, E. R. Schreiter, Labeling of active neural

circuits in vivo with designed calcium integrators.

Science

347

755

–

760 (2015).

50. H. H. Yang, F. St-Pierre, Genetically encoded voltage indicators: Opportunities and

challenges.

J. Neurosci.

, 9977

–

9989 (2016).

51. E. S. Boyden, F. Zhang, E. Bamberg, G. Nagel, K. Deisseroth, Millisecond-timescale,

genetically targeted optical control of neural activity.

Nat. Neurosci.

, 1263

–

1268 (2005).

52. D. G. Ouzounov, T. Wang, M. Wang, D. D. Feng, N. G. Horton, J. C. Cruz-Hernández,

Y.-T. Cheng, J. Reimer, A. S. Tolias, N. Nishimura, C. Xu, In vivo three-photon imaging of

activity of GCaMP6-labeled neurons deep in intact mouse brain.

Nat. Methods

388

–

390 (2017).

53. J. P. Rickgauer, D. W. Tank, Two-photon excitation of channelrhodopsin-2 at saturation.

Proc. Natl. Acad. Sci. U.S.A.

106

, 15025

–

15030 (2009).

54. B. K. Andrasfalvy, B. V. Zemelman, J. Tang, A. Vaziri, Two-photon single-cell optogenetic

control of neuronal activity by sculpted light.

Proc. Natl. Acad. Sci. U.S.A.

107

11981

–

11986 (2010).

55. E. Papagiakoumou, F. Anselmi, A. Bègue, V. de Sars, J. Glückstad, E. Y. Isacoff, V. Emiliani,

Scanless two-photon excitation of channelrhodopsin-2.

Nat. Methods

,848

–

854 (2010).

56. R. Prakash, O. Yizhar, B. Grewe, C. Ramakrishnan, N. Wang, I. Goshen, A. M. Packer,

D. S. Peterka, R. Yuste, M. J. Schnitzer, K. Deisseroth, Two-photon optogenetic toolbox for fast

inhibition, excitation and bistable modulation.

Nat. Methods

, 1171

–

1179 (2012).

57. A.M.Packer,D.S.Peterka,J.J.Hirtz,R.Prakash,K.Deisseroth,R.Yuste,Two-photon

optogenetics of dendritic spines and neural circuits.

Nat. Methods

, 1202

–

1205 (2012).

58. F.Zhang,V.Gradinaru,A.R.Adamantidis,R.Durand,R.D.Airan,L.deLecea,K.Deisseroth,

Optogenetic interrogation of neural circuits: Technology for probing mammalian brain

structures.

Nat. Protoc.

, 439

–

456 (2010).

59. I. Yamaguchi, T. Zhang, Phase-shifting digital holography.

Opt. Lett.

, 1268

–

1270 (1997).

60. H. Ruan, M. Jang, B. Judkewitz, C. Yang, Iterative time-reversed ultrasonically encoded

light focusing in backscattering mode.

Sci. Rep.

, 7156 (2014).

61. K. Si, R. Fiolka, M. Cui, Breaking the spatial resolution barrier via iterative sound-light

interaction in deep tissue microscopy.

Sci. Rep.

, 748 (2012).

62. Y. Suzuki, J. W. Tay, Q. Yang, L. V. Wang, Continuous scanning of a time-reversed

ultrasonically encoded optical focus by reflection-mode digital phase conjugation.

Opt. Lett.

, 3441

–

3444 (2014).

63. C. Xiao, J. R. Cho, C. Zhou, J. B. Treweek, K. Chan, S. L. McKinney, B. Yang, V. Gradinaru,

Cholinergic mesopontine signals govern locomotion and reward through dissociable

midbrain pathways.

Neuron

, 333

–

347 (2016).

64. J. Ryan Cho, J. B. Treweek, J. Elliott Robinson, C. Xiao, L. R. Bremner, A. Greenbaum,

V. Gradinaru, Dorsal raphe dopamine neurons modulate arousal and promote

wakefulness by salient stimuli.

Neuron

1205

–

1219.e8 (2017).

65. C. K. Kim, S. J. Yang, N. Pichamoorthy, N. P. Young, I. Kauvar, J. H. Jennings, T. N. Lerner,

A. Berndt, S. Y. Lee, C. Ramakrishnan, T. J. Davidson, M. Inoue, H. Bito, K. Deisseroth,

Simultaneous fast measurement of circuit dynamics at multiple sites across the

mammalian brain.

Nat. Methods

, 325

–

328 (2016).

66. J. Y. Lin, P. M. Knutsen, A. Muller, D. Kleinfeld, R. Y. Tsien, ReaChR: A red-shifted variant of

channelrhodopsin enables deep transcranial optogenetic excitation.

Nat. Neurosci.

1499

–

1508 (2013).

67. M. Z. Lin, M. J. Schnitzer, Genetically encoded indicators of neuronal activity.

Nat. Neurosci.

, 1142

–

1153 (2016).

68. M. Jang, C. Yang, I. M. Vellekoop, Optical phase conjugation with less than a photon per

degree of freedom.

Phys. Rev. Lett.

118

, 093902 (2017).

69. M. Jang, H. Ruan, B. Judkewitz, C. Yang, Model for estimating the penetration depth limit

of the time-reversed ultrasonically encoded optical focusing technique.

Opt. Express

5787

–

5807 (2014).

70. A. S. Chuong, M. L. Miri, V. Busskamp, G. A. C. Matthews, L. C. Acker, A. T. Sørensen,

A. Young, N. C. Klapoetke, M. A. Henninger, S. B. Kodandaramaiah, M. Ogawa,

S. B. Ramanlal, R. C. Bandler, B. D. Allen, C. R. Forest, B. Y. Chow, X. Han, Y. Lin, K. M. Tye,

B. Roska, J. A. Cardin, E. S. Boyden, Noninvasive optical inhibition with a red-shifted

microbial rhodopsin.

Nat. Neurosci.

, 1123

–

1129 (2014).

71. A. A. Kaberniuk, A. A. Shemetov, V. V. Verkhusha, A bacterial phytochrome-based

optogenetic system controllable with near-infrared light.

Nat. Methods

, 591

–

597

(2016).

72. J. Brake, M. Jang, C. Yang, Analyzing the relationship between decorrelation time and

tissue thickness in acute rat brain slices using multispeckle diffusing wave spectroscopy.

J. Opt. Soc. Am. A

, 270

–

275 (2016).

73. A. Lev, B. Sfez, In vivo demonstration of the ultrasound-modulated light technique.

J. Opt. Soc. Am. A

, 2347

–

2354 (2003).

74. M. M. Qureshi, J. Brake, H.-J. Jeon, H. Ruan, Y. Liu, A. M. Safi, T. J. Eom, C. Yang, E. Chung,

In vivo study of optical speckle decorrelation time across depths in the mouse brain.

Biomed. Opt. Express

, 4855

–

4864 (2017).

SCIENCE ADVANCES

RESEARCH ARTICLE

Ruan

et al

Sci. Adv.

2017;

:eaao5520 8 December 2017

8of9

on December 12, 2017

http://advances.sciencemag.org/

Downloaded from

75. T. Laforest, A. Verdant, A. Dupret, S. Gigan, F. Ramaz, G. Tessier, Co-integration of a smart

CMOS image sensor and a spatial light modulator for real-time optical phase modulation.

Proc. SPIE

9022

, 90220N (2014).

76. M. Jang, H. Ruan, H. Zhou, B. Judkewitz, C. Yang, Method for auto-alignment of digital

optical phase conjugation systems based on digital propagation.

Opt. Express

14054

–

14071 (2014).

77. H. Ruan, M. L. Mather, S. P. Morgan, Pulsed ultrasound modulated optical tomography

with harmonic lock-in holography detection.

J. Opt. Soc. Am. A

, 1409

–

1416

(2013).

Acknowledgments:

We would like to thank B. Yang, L. Bremner, A. Shibukawa, and H. Deng

for their assistance and helpful discussions.

Funding:

We would like to acknowledge

support from the NIH (DP2OD007307 to C.Y., U01NS090577 to C.Y. and V.G., and F31EB021153

to J.B.), the Gwangju Institute of Science and Technology

–

California Institute of Technology

(Caltech) Collaborative Research Program (CG2012 to C.Y.), the Children

’

s Tumor Foundation

(2016-01-006 to J.E.R.), the Donna and Benjamin M. Rosen Bioengineering Center (to J.B.),

the Heritage Medical Research Institute (to V.G.), and the Tianqiao and Chrissy Chen Institute for

Neuroscience at Caltech (to V.G.).

Author contributions:

H.R. and J.B. contributed equally to

the work. H.R. designed the experimental setup. H.R., J.B., M.J., and Y.L. conducted the optical

experiments. J.E.R., C.X., and C.Z. prepared the biological samples. J.E.R., J.B., C.X., and C.Z.

conducted the electrophysiological recordings. H.R., J.B., Y.L., and J.E.R. analyzed the experimental

data. V.G. and C.Y. supervised the project. All authors contributed to the manuscript preparation.

Competing interests:

C.Y. is an author on a patent related to this work (publication no.

US9313423 B2, filed on 27 March 2013). The authors declare that they have no other competing

interests.

Data and materials availability:

All data needed to evaluate the conclusions in the

paper are present in the paper and/or the Supplementary Materials. Additional data related to this

paper may be requested from the authors.

Submitted 2 August 2017

Accepted 8 November 2017

Published 8 December 2017

10.1126/sciadv.aao5520

Citation:

H. Ruan, J. Brake, J. E. Robinson, Y. Liu, M. Jang, C. Xiao, C. Zhou, V. Gradinaru, C. Yang,

Deep tissue optical focusing and optogenetic modulation with time-reversed ultrasonically

encoded light.

Sci. Adv.

, eaao5520 (2017).

SCIENCE ADVANCES

RESEARCH ARTICLE

Ruan

et al

Sci. Adv.

2017;

:eaao5520 8 December 2017

9of9

on December 12, 2017

http://advances.sciencemag.org/

Downloaded from

encoded light

Deep tissue optical focusing and optogenetic modulation with time-reversed ultrasonically

Changhuei Yang

Haowen Ruan, Joshua Brake, J. Elliott Robinson, Yan Liu, Mooseok Jang, Cheng Xiao, Chunyi Zhou, Viviana Gradinaru and

DOI: 10.1126/sciadv.aao5520

(12), eaao5520.

Sci Adv

ARTICLE TOOLS

http://advances.sciencemag.org/content/3/12/eaao5520

MATERIALS

SUPPLEMENTARY

http://advances.sciencemag.org/content/suppl/2017/12/04/3.12.eaao5520.DC1

REFERENCES

http://advances.sciencemag.org/content/3/12/eaao5520#BIBL

This article cites 77 articles, 6 of which you can access for free

PERMISSIONS

http://www.sciencemag.org/help/reprints-and-permissions

Use of this article is subject to the

registered trademark of AAAS.

is a

Science Advances

Association for the Advancement of Science. No claim to original U.S. Government Works. The title

(ISSN 2375-2548) is published by the American Association for the Advancement of Science, 1200 New

Science Advances

on December 12, 2017

http://advances.sciencemag.org/

Downloaded from