untitled - PhysRevLett.112.014302.pdf

Photoimprint Photoacoustic Microscopy for Three-Dimensional Label-Free

Subdiffraction Imaging

Junjie Yao, Lidai Wang, Chiye Li, Chi Zhang, and Lihong V. Wang

Optical Imaging Laboratory, Department of Biomedical Engineering, Washington University in St. Louis,

St. Louis, Missouri 63130, USA

(Received 19 September 2013; published 10 January 2014)

Subdiffraction optical microscopy allows the imaging of cellular and subcellular structures with a res-

olution finer than the diffraction limit. Here, combining the absorption-based photoacoustic effect and

intensity-dependent photobleaching effect, we demonstrate a simple method for subdiffraction photoacous-

tic imaging of both fluorescent and nonfluorescent samples. Our method is based on a double-excitation

process, where the first excitation pulse partially and inhomogeneously bleaches the molecules in the dif-

fraction-limited excitation volume, thus biasing the signal contributions from a second excitation pulse

striking the same region. The differential signal between the two excitations preserves the signal contri-

bution mostly from the center of the excitation volume, and dramatically sharpens the lateral resolution.

Moreover, due to the nonlinear nature of the signal, our method offers an inherent optical sectioning

capability, which is lacking in conventional photoacoustic microscopy. By scanning the excitation beam,

we performed three-dimensional subdiffraction imaging of varied fluorescent and nonfluorescent species.

As any molecules have absorption, this technique has the potential to enable label-free subdiffraction im-

aging, and can be transferred to other optical imaging modalities or combined with other subdiffraction

methods.

DOI:

10.1103/PhysRevLett.112.014302

PACS numbers: 43.35.Ud, 42.25.Bs, 42.25.Fx, 87.64.M-

In recent years, by breaking the diffraction limit, subdif-

fraction optical microcopy has revolutionized fundamental

biological studies. Generally speaking, subdiffraction tech-

niques fall into two broad categories: so-called

“

pattern

excitation

”

approaches and single-molecule localization

approaches

[1]

. In the method we describe here, the reso-

lution enhancement is based on the excitation nonlinearity

of the photobleaching effect, a common phenomenon in

optical imaging which is otherwise regarded as harmful

[2,3]

. The photobleaching effect depends strongly on the

excitation intensity for both fluorescent and nonfluorescent

species, which enables subdiffraction imaging by spatially

trimming the excitation volume to a subdiffraction size

–

. Since all molecules are optically absorbing at

selected wavelengths, photoacoustic (PA) imaging, which

acoustically probes optical absorption contrast in biological

tissue, can potentially image all molecules, endogenous and

exogenous

[7]

. Therefore, the combination of the photo-

bleaching effect and photoacoustic imaging can potentially

achieve subdiffraction imaging over a wide range of species.

Photoacoustic imaging is based on the photoacoustic

effect. The principle of photobleaching-based photoimprint

subdiffraction PA microscopy (PI-PAM) is illustrated in

Fig.

1(a)

. When a Gaussian-shape diffraction-limited exci-

tation spot strikes on densely distributed absorbers, the gen-

erated PA signal is a summation of the contributions from

all absorbers inside the excitation spot [Fig.

1(a)

, left

panel]. After the first excitation, the absorbers inside the

excitation spot are inhomogeneously bleached, depending

on the local excitation intensity [Fig.

1(a)

, middle panel].

Therefore, the reduction of absorption in the center of the

excitation spot is greater than that in the periphery. As a

result, when the second pulse excites the same region,

the center portion contributes less to the second PA signal

than the periphery. The difference between the two PA sig-

nals not only reflects the excitation intensity profile, but

also incorporates the absorption reduction distribution

[Fig.

1(a)

, right panel], which therefore sharpens the center

of the focus. This concept of enhancement in the lateral

resolution is elucidated in Fig.

1(b)

. In summary, while

each PA signal is linear to the excitation intensity, the dif-

ferential signal is nonlinear to the excitation intensity. This

is the physical basis of our method.

The contrast of the PI-PAM comes from the differential

signal between two adjacent frames, expressed as (see

Supplemental Material

[8]

, Note 1, for more derivation)

−

1

−

∝

0

1

;

(1)

where

is the signal amplitude detected by the ultrasonic

transducer with the

th excitation,

is the Grueneisen coef-

ficient,

is the percentage of the absorbed photon energy

that is converted into heat,

0

is the initial number of mol-

ecules,

is the excitation intensity, and

is the power

dependence of the photobleaching rate on the excitation

intensity. Equation

(1)

indicates that, on the one hand,

the PI-PAM signal is linear to the optical absorption, which

maintains its functional imaging capability, such as oxygen

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saturation measurement. On the other hand, the PI-PAM

signal is nonlinear to the excitation intensity, which enables

the subdiffraction imaging capability.

If the excitation profile can be approximated by a

Gaussian function, we obtain the full-width at half-

maximum (FWHM) of the lateral point spread function

(PSF) of the imaging system as (see Supplemental

Material

[8]

, Note 2, for more detailed derivation)

FWHM

ffiffiffiffiffiffiffiffiffiffiffiffi

2

2

1

≈

0

51

ffiffiffiffiffiffiffiffiffiffiffi

1

0

;

(2)

where

is the radial distance from the center of the Airy

disk,

is the Gaussian width of the excitation beam where

the beam intensity drops to

1

2

of its center value,

0

is the

excitation wavelength and NA is the numerical aperture of

the objective. Equation

(2)

indicates that the effective PSF

[Fig.

1(b)

, right panel] of the system is sharper than the

initial diffraction-limited excitation PSF [Fig.

1(b)

, left

panel] by a factor of

ffiffiffiffiffiffiffiffiffiffiffi

1

Similarly, the axial resolution of PI-PAM for point tar-

gets is given by the FWHM of the axial PSF as (see

Supplemental Material

[8]

, Note 3, for more detailed

derivation)

FWHM

point

2

ffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi

2

1

1

−

1

1

8

ffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi

2

1

1

−

1

0

2

;

(3)

where

is the axial distance from the focal plane and

the Rayleigh range of the Gaussian beam. Equation

(3)

shows that, by sharpening the optical focal zone,

PI-PAM provides an axial resolution improvement by a fac-

tor of

1

ffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi

2

1

1

−

1

over conventional PAM for point

targets.

For planar (or wide) targets, conventional PA micros-

copy (PAM) lacks sectioning capability because its axial

PSF is constant. By contrast, for PI-PAM, the optical sec-

tioning capability can be expressed as the FWHM of the

axial PSF:

FWHM

planar

2

ffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi

2

1

−

1

1

8

ffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi

2

1

−

1

0

2

. (4)

Equation

(4)

shows that the optical sectioning capability

of PI-PAM also depends on the intensity power dependence

of the photobleaching rate.

The photoimprint PA microscopy system is illustrated in

Fig.

1(c)

(for more details, see Supplemental Material

[8]

methods and Fig. 1). Briefly stated, an OPO (optical para-

metric oscillator) laser with a wavelength tuning range of

210

–

2600 nm serves as the excitation source. After spatial

filtering, the beam is focused into the sample by a 1.4 NA

(numerical aperture) oil-immersion objective. The photo-

acoustic waves are detected by a focused ultrasonic trans-

ducer (customized with 40 MHz central frequency, 80%

bandwidth, and 0.50 NA), which is placed confocally with

the objective and coupled by water. A customized

ABVIEW

program synchronizes the entire system.

PI-PAM imaging can be performed point by point (

-scan

mode), cross section by cross section (

-scan mode), or

volume by volume (

-scan mode). If not otherwise men-

tioned,

-scan mode was used for the experiments in this

Letter. The second pulse is a separate event, which follows

the first pulse that partially bleaches the absorbers. To

ensure that the PA signals from two consecutive pulses

do not temporally overlap, the time interval between the

two pulses should be at least a few microseconds. In prac-

tice, we make sure that the two pulses are more than 1 ms

FIG. 1 (color online). Photoimprint photoacoustic microscopy

(PI-PAM). (a) Principle of PI-PAM. The differential signal be-

tween before- (left panel) and after-bleaching (middle panel) im-

ages results in a smaller effective excitation size, as shown by the

dashed circle in the right panel. (b) Illustration of the lateral res-

olution enhancement by PI-PAM. The effective PSF is the prod-

uct of the excitation PSF and the photobleaching profile.

Beam sampler (BS), neutral density filter (ND), photodiode

(PD), optical parametric oscillator (OPO), and ultrasonic trans-

ducer (UT).

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apart. To achieve sufficient detection sensitivity, typically

more than 200 frames were acquired for each experiment.

To improve the signal-to-noise ratio (SNR) in practice,

we took the difference between frames that are half of

the total number of frames apart, and then averaged all

the differential images (see Supplemental Material

[8]

Note 4, for more detailed derivation). Assume that the total

number of frames is

, the true signal difference between

adjacentframesis

,andthestandarddeviationofthewhite

noise is

. Approximately, averaging of the differential sig-

nals between frames that are

2

apart yields a SNR of

SNR

2

4

ffiffiffiffiffi

(5)

The SNR ratio between PI-PAM and conventional PAM

is given by

SNR

PI-PAM

SNR

C-PAM

2

2

−

(6)

Equation

(6)

suggests that the SNR advantage of conven-

tional PAM decreases as the bleached percentage

increases. The SNR of PI-PAM can be improved by

increasing the bleaching percentage with either more imag-

ing frames or stronger excitation intensity.

Using this setup, we first studied the photobleaching

characteristics of several biological and nonbiological sam-

ples commonly used in PA imaging, including red blood

cells, B16 melanoma cells, black ink, gold nanoparticles,

and graphite (Supplemental Material

[8]

, Fig. 2). An expo-

nential-decay model was used to fit for the photobleaching

rate at different excitation intensities. The power depend-

ence of each material on the excitation intensity determines

the expected resolution improvement. Among the test

materials, gold nanoparticles have the highest power

dependence of 4.0, which is expected to improve the lateral

resolution to

∼

87

nm.

The lateral resolution enhancement of PI-PAM over con-

ventional PA microscopy is presented in Fig.

. A sharp

blade edge coated with hemoglobin was imaged at

532 nm, and the corresponding line spread function was

fitted to compute the lateral resolution [Fig.

2(a)

]. With

an objective NA of 1.4, conventional PAM achieved a lat-

eral resolution of 200 nm, in agreement with the diffraction

limit. By contrast, PI-PAM achieved a subdiffraction lateral

resolution of 120 nm, a 1.7-fold improvement. Since the

photobleaching of hemoglobin has a power dependence

of 2.2 on the excitation intensity, the measured resolution

is close to the expected value of 112 nm. To demonstrate

the imaging performance, conventional PAM and PI-PAM

were applied to image a monolayer of 200-nm-diameter

densely packed red-dyed microspheres [Fig.

2(b)

]. The

red dye

’

s photobleaching property was similar to that of

hemoglobin. The images show that the microspheres

were much better resolved by PI-PAM, which confirmed

the lateral resolution improvement.

We further demonstrate the subdiffraction imaging

capability of PI-PAM on both nonbiological and biological

samples (Fig.

). We compared conventional PAM and

PI-PAM images of 150-nm diameter gold nanoparticles

[Fig.

3(a)

and movie S1 in the Supplemental Material

[8]

]. Two closely located nanoparticles, 270 nm apart, were

clearly resolved by PI-PAM but barely resolved by conven-

tional PAM [Fig.

3(a)

, insets]. The line profile across one

nanoparticle shows a full width at half maximum of 170 nm

[Fig.

3(a)

, right]. Because the line profile is a convolution

of the effective PSF of the system with the particle absorp-

tion profile, deconvolution of the line profile using the

Gaussian approximation results in an effective lateral

resolution of 80 nm for PI-PAM, consistent with the

expected value. We also imaged a fixed B16 melanoma cell

[Fig.

3(b)

and movie S2

[8]

]. The PI-PAM image is clearly

superior in resolution. Small dendrites in the melanoma cell

-600

-400

-200

200

400

600

-0.5

0.0

0.5

1.0

1.5

x

(nm)

Normalized PA amplitude

(a)

(b)

PI-PAM

Conventional PAM

Normalized PA amplitude

500 nm

x

y

-400

400

0.0

0.5

1.0

120 n

200 nm

150 nm

PI-PAM

Conventional PAM

FIG. 2 (color online). Lateral resolution enhancement by

PI-PAM. (a) Edge spread function of conventional PAM and

PI-PAM, using a sharp blade edge coated with hemoglobin.

(b) A cluster of red-dyed microspheres with nominal diameters

of 200 nm, as imaged by conventional PAM (left) and

PI-PAM (right). The dashed boxes in the top row were magnified

and shown in the bottom row.

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that are less than 100 nm apart can be clearly separated by

PI-PAM [Fig.

3(b)

, insets].

The other important advantage offered by PI-PAM is its

optical sectioning capability, providing axial resolution for

not only point targets but also targets that are large relative

to the lateral resolution. As shown by Eq.

(4)

, the axial res-

olution for planar targets originates from the nonuniform

light intensity along the axial beam axis. The light intensity

is stronger at the focal plane. Therefore, the absorbers at the

focal plane are bleached more than the absorber farther

away from the focal plane. The difference of the PA signals

before and after bleaching is thus axial-position dependent.

This dependence results in the axial resolution for planar

targets. A simple illustration is shown in Fig.

4(a)

. Like

the lateral resolution enhancement, the sectioning strength

of PI-PAM is determined by the power dependence of the

photobleaching rate on the excitation intensity. For exam-

ple, with a linear dependence, PI-PAM should achieve a

sectioning capability comparable with that of confocal

microscopy. We first quantified the sectioning by using a

layer of dried hemoglobin (

∼

150

nm in thickness)

[Fig.

4(b)

and Supplemental Material, Fig. 4

[8]

]. While

the signal strength in conventional PAM decayed very

slowly, the signal strength in PI-PAM sharply dropped

as the sample moved out of focus. The FWHM of the signal

profile in PI-PAM suggests a sectioning capability of

370 nm after deconvolution with the absorption profile

of the sample, close to the expected value (320 nm).

Accordingly,a single layer ofred bloodcells fixedon a cover

glass was imaged at varying axial positions [Fig.

4(c)

and

movie S3 in the Supplemental Material

[8]

]. The stack of

themaximum-amplitudeprojectionsontheorthogonalplane

clearly shows the optical sectioning capability of PI-PAM

[Fig.

4(c)

Finally, we applied PI-PAM to subdiffraction imaging of

live rose petal epidermal cells at 570 nm. Pink anthocyanins

in cell vacuoles provided the absorption contrast for PA im-

aging, and they are also commonly used for fluorescence

imaging (Supplemental Material, Fig. 4

[8]

). In the PA

images, densely packed epidermal cells can be clearly

resolved. The cell vacuoles are bright, while the cell walls

and cell nuclei are dark [Fig.

4(d)

]. By

scanning the sam-

ple with a step size of

0

5

m, PI-PAM was performed to

FIG. 3 (color online). PI-PAMimagingofgoldnanoparticlesand

melanoma cells with enhanced lateral resolution. (a)

–

(b) 150-nm

diameter gold nanoparticles (a), and a B16 melanoma cell (b) im-

aged by conventional PAM (left column) and PI-PAM (middle

column). The normalized signal amplitude profiles along the

dashed lines are shown in the right column. Insets show magni-

fied views of the marked areas. The scale bars for the insets are

200 nm.

-2

-1

0.0

0.2

0.4

0.6

0.8

1.0

z

(

Normalized PA amplitude

~400 nm

Before bleaching

After bleaching

PI-PAM

(a)

(b)

(c)

x

z

Normalized laser

intensity

Absorbers

PI-PAM

Conventional PAM

x

z

10 μm

Normalized

PA amplitude

(d)

10 μm

Conventional PAM

PI-PAM

Normalized PA

amplitude

Conventional

PAM

PI-PAM

(f)

(e)

20 μm

z

=0μm

z

=2μm

z

=4μm

z

=0μm

z

=2μm

z

=4μm

20 μm

FIG. 4 (color online). Optical sectioning capability of PI-PAM.

(a) Illustration of the optical sectioning capability provided by

PI-PAM. The differential signal of the before- (left) and after-

bleaching (middle) images largely removes the out-of-focus

contributions. (b) The axial point spread function measured on

a thin-layer hemoglobin sample. (c) Orthogonal projections of

single layer red blood cells imaged by conventional PAM and

PI-PAM at varying axial positions. (d)

–

(f) Subdiffraction imaging

of live rose petal epidermal cells at 570 nm. The sample was

scanned with a step size of

0

5

m. Maximum-amplitude projec-

tions (d) were extracted from the

-stacked images (e)

–

(f).

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section the cells at different depths [Figs.

4(e)

and

4(f)

and movie S4 in the Supplemental Material

[8]

]. With

scanning, PI-PAM has achieved three-dimensional subdif-

fraction resolution.

In summary, we have demonstrated three-dimensional

subdiffraction imaging of both fluorescent and nonfluores-

cent samples by spatially trimming the excitation volume

via the photobleaching effect. We note that PI-PAM can

be applied to both fluorescent and nonfluorescent species

[9]

. When fluorescence is present, it may change the

bleaching rate power dependence on the excitation inten-

sity, and thus the resolution improvement. Fortunately,

most fluorescent species have a quite low quantum yield

(typically less than 10%); therefore, the nonfluorescent

bleaching effect still dominates the imaging process.

It is necessary to discuss the bleaching effect on the rel-

ative feature contrast between the

“

bleached region

”

and

“

nonbleached region.

”

First, approximately, only the

absorbers within the optical focal plane are bleached.

When the optical focal plane is moved to other depths,

the same bleaching process is repeated and thus the final

image contrast between different depths should not be

changed. Second, the distortion between the bleached

and nonbleached regions can be minimized by reducing

the frame numbers or the excitation light intensity.

However, this results in reduced SNR as well. In practice,

a balance between sample distortion and SNR should be

carefully determined. If permanent photobleaching is a

concern in the imaging process, the dynamics of reversibly

photoswitchable chromophores can be used instead. While

the photoswitchable chromophores can be cycled between

bright and dark states, the switching rate strongly depends

on the switching light intensity

[10,11]

, which provides a

zero-loss alternative for PI-PAM.

The same principle of PI-PAM can be readily transferred

to fluorescence microscopy, the major imaging tool cur-

rently used in biological studies. Because the photobleach-

ing effect occurs in the excitation phase of the imaging

process, our method can be applied to confocal microscopy

and two-photon microscopy. Compared with other optical

imaging modalities, PA detection is not necessarily supe-

rior, because each modality has its own unique advantages.

Our main purpose is to introduce a subdiffraction imaging

method to both PA and other optical imaging communities.

Compared with fluorescence microscopy, photoacoustic

microscopy can image nonfluorescent species, and can

potentially penetrate deeper due to the one-way optical

scattering. However, compared with optical imaging where

high sensitivity detectors are used, PA imaging suffers from

the relatively low SNR due to the less sensitive ultrasonic

transducers. To further improve the PA detection sensitiv-

ity, more work is necessary to develop optimum illumina-

tion and state-of-the-art acoustic detectors

[12]

. We expect

that realizable improvements to the acoustic detector may

enable single molecule detection sensitivity, comparable to

that of optical detection

[12]

The authors appreciate support by Professor James

Ballard and useful discussions with Konstantin Maslov,

Amos Danielli, and Arie Krumholz. This research was sup-

ported by the National Institutes of Health Grants No. DP1

EB016986, No. R01 EB008085, No. R01 CA134539,

No. U54 CA136398, No. R01 CA157277, and No. R01

CA159959.

lhwang@wustl.edu

[1] B. Huang, M. Bates, and X. W. Zhuang,

Annu. Rev.

Biochem.

, 993 (2009)

[2] K. Jacobson, E. Elson, D. Koppel, and W. Webb,

Nature

(London)

295

, 283 (1982)

[3] I. Rasnik, S. A. McKinney, and T. Ha,

Nat. Methods

, 891

(2006)

[4] L. Gao, L. Wang, C. Li, A. Garcia-Uribe, and L. V. Wang,

J. Biophotonics

, 543 (2013)

[5] T. S. Chen, S.-Q Zeng, Q.-M. Luo, Z.-H. Zhang, and

W. Zhou,

Biochem. Biophys. Res. Commun.

291

, 1272

(2002)

[6] G. H. Patterson and D. W. Piston,

Biophys. J.

, 2159

(2000)

[7] L. H. V. Wang and S. Hu,

Science

335

, 1458 (2012)

[8] See Supplemental Material at

http://link.aps.org/

supplemental/10.1103/PhysRevLett.112.014302

for de-

tailed derivations of the equations, as well as other support-

ing notes, methods, figures, and videos.

[9] H. F. Launer,

Nature (London)

218

, 160 (1968)

[10] M. Bates, B. Huang, B., G. T. Dempsey, and X. Zhuang,

Science

317

, 1749 (2007)

[11] G. H. Patterson, L.-M. Ting, Y. Wang, J. S. Condeelis, and

V. V. Verkhusha,

Nat. Methods

, 771 (2011)

[12] A. M. Winkler, K. Maslov, and L. V. Wang,

J. Biomed. Opt.

, 097003 (2013)

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