OPTICS
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Deep tissue optical focusing and optogenetic
modulation with time-reversed ultrasonically
encoded light
Haowen Ruan,
1
* Joshua Brake,
1
* J. Elliott Robinson,
2
Yan Liu,
1
Mooseok Jang,
1
†
Cheng Xiao,
2
‡
Chunyi Zhou,
2
‡
Viviana Gradinaru,
2
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
l
≈
26
m
m) (
1
,
2
), 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 (
3
–
10
) 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 (
11
). 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 (
11
) and diffuse optical tomography (
12
), 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 (
13
). 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
(
3
,
8
,
14
,
15
), transmission matrix measurement (
9
,
10
,
16
–
19
), or op-
tical time reversal (optical phase conjugation) (
13
,
20
–
24
). 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 (
25
–
30
). 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 (
25
–
27
,
30
). When used with a guide-
star (
5
), 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
(
31
,
32
), nonlinear optical (
22
,
33
–
36
), kinetic (
37
,
38
), photoacoustic
(
39
–
42
), ultrasonic (
43
–
46
), magnetic (
47
), and microbubble (
48
)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
m
m)
(
44
–
46
).
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
1
Department of Electrical Engineering, California Institute of Technology, 1200 East
California Boulevard, Pasadena, CA 91125, USA.
2
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
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genetically encoded neural activity indicators (
49
,
50
)andcontrolling
activity via optogenetic actuators (
51
). Although neurophotonic tech-
niques that use multiphoton excitation (
52
–
57
) and adaptive optics
(
33
,
34
,
36
) 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 (
58
). 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
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
(
21
,
22
) 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 (
59
). 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 (
21
,
22
), which, compared
with analog TRUE fo
cusing systems (
44
), allowed for measured wave-
fronts to be played back at a light intensity far greater than that of the
measured wave (
45
).
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 (
32
)
h
phase
‐
only
¼
p
4
ð
N
1
Þ
M
þ
1
ð
1
Þ
where
N
is the number of optical modes controlled by the SLM and
M
is the number of optical modes (speckle grains) within the ultrasound
A
B
C
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
) A DIC microscope was included for
neuron visualization during patch clamping. (
B
) The TRUE focusing system allowed
light to be sharply focused through the brain slice. (
C
)Aclose-upimageoftheTRUE
focus on a patched neuron. Scale bars, 20
m
min(A)and50
m
min(B).
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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 (
60
–
62
), 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 (
61
), 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
m) that contained the medial
prefrontal cortex (mPFC)
from C57Bl/6J mice using a vibrating microtome as previously de-
scribed (
63
,
64
). 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
m-thick slice, the lateral width of the focus profile was signif-
icantly increased from the diffraction-limited focus size of ~1
m
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
–
m
m
2
field of view
in the 1000-
m
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
m
m,
and the calibrated waveform duration (
−
6 dB) was 37.4 ns, correspond-
ing to a pulse length of 55.3
m
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 (
60
) to achieve a
focus with an average FWHM spot size of 27.4
m
m across tissue thick-
nesses from 500 to 2000
m
m (see method S1 for calculation). In con-
trast, the FWHM of the conventional focus broadened from ~350
m
mat
a slice thickness of 500
m
m to approximately 2100
m
matathicknessof
2000
m
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
m thick, which, unlike those formed by conventional focusing,
do not significantly broaden with increased sample thickness.
0.5
1
1.5
2
10
1
10
2
10
3
10
4
Focus FWHM versus thickness
Tissue thickness (mm)
Focus FWHM (
μ
m)
Conventional
TRUE
C
D
B
A1
A2
1
0
1
0
Camera
SLM
Beam splitter
Beam splitter
Camer
a
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.
(
A
) The recording (A1) and playback (A2) procedures used to focus light through the slice onto
its top surface with TRUE focusing. (
B
) 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. (
C
) Images of the
conventional and TRUE focus profile through living brain tissue slices (500, 800, 1000, 1500, and 2000
m
m thick). (
D
) 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
m
m.
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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 (
51
). 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) (
65
). 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 (
66
), bReaChES displayed
a nonlinear increase in photocurre
nt response that saturated at an in-
tensity of approximately 10 mW/mm
2
(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
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
m
mat532nm)(
2
). 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% (
n
= 6) in 800-
m
m-thick brain
slices, which was similar in magn
itude to the enhancement observed
in 300- and 500-
m
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
6
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
h
of ~1 × 10
4
.Thisexper-
imental performance means the DOPC system could effectively control
N
~1×10
4
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
10
Light intensity (mW/mm
2
)
0
0.2
0.4
0.6
1.0
0.8
Normalized photocurren
t
bReaChES photocurren
t
n
= 10
10
1
10
0
10
1
10
2
CD
AB
2
Fig. 3. Experiment design, opsin characterization, and demonstration of photo-
current and firing modulation via TRUE focusing.
(
A
) An AAV vector was used to
stereotaxically deliver the bReaChES transgene to the mPFC. (
B
) Characterization
of normalized photocurrent response versus light intensity. The average maxi-
mum photocurrent across the 10 cells studied was 1047 pA. (
C
) Diagram illustrat-
ing the experimental scheme used to demonstrate the ability of TRUE focusing to
elicit action potentials through 800-
m
m-thick living mouse brain tissue. (
D
) Repre-
sentative traces demonstrating elicited photocurrent and membrane voltage changes
achieved with and without TRUE focusing.
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diameter of ~27
m
m, the number of modes
M
inside the focus was
~1 × 10
4
, 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
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
m
minthe
x
and
y
dimensions. In contrast, the FWHMs for the
TRUE focusing scan were 99 and 71
m
minthe
x
and
y
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 (
58
), 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
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 (
67
), 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.
x
y
x
z
B
A
CD
800 μm
Electrode
Lens
Brain slice
Photocurrent
Conventional
TRUE
0
0.2
0.4
0.6
0.8
1
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 (
A
)andTRUEfocusing(
B
). The normalized
photocurrent enhancement as a function of lateral focal scanning position for
conventional (
C
) and TRUE focusing (
D
). Scale bar, 100
m
m.
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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) (
68
), the presence of shot noise due to the
much higher unmodulated light intensity will fundamentally limit
the penetration depth (
24
,
69
). Additional guidestar aids, such as micro-
bubbles, can help improve the tagging efficiency significantly (
48
)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
N
, the number of controllable
modes, or decrease
M
, the number of optical modes within the ultra-
sound focus. One way to reduce
M
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 (
70
,
71
). A more feasible avenue to improve the TRUE focus
contrast is to increase the number of controllable optical modes,
N
.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
N
by 10 times will
result in a focus intensity
–
to
–
background ratio
h
≈
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 (
72
),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. (
26
,
73
,
74
). 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 (
25
,
27
), 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 (
75
). 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 (
76
), 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
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
(
63
,
64
). 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
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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 (
59
) 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,
p
/2,
p
,
and 3
p
/2. Four intensity maps (
I
0
,
I
p
/2
,
I
p
,and
I
3
p
/2
) corresponding to
each phase of the reference beam were recorded, and the phase map of
the ultrasonically tagged light was calculated as
f
= Arg[(
I
p
/2
−
I
3
p
/2
)+
i
(
I
0
−
I
p
)], 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 (
77
), 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
m-thick acute
brain slices.
fig. S5. Electrophysiological photocurrent and membrane voltage traces comparing ultrasound
on and off conditions.
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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.
3
, eaao5520 (2017).
SCIENCE ADVANCES
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RESEARCH ARTICLE
Ruan
et al
.,
Sci. Adv.
2017;
3
:eaao5520 8 December 2017
9of9
on December 12, 2017
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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.
3
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