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Single-cell photoacoustic

thermometry

Liang Gao, Lidai Wang, Chiye Li, Yan Liu, Haixin Ke, et

al.

Liang Gao, Lidai Wang, Chiye Li, Yan Liu, Haixin Ke, Chi Zhang, Lihong V.

Wang , "Single-cell photoacoustic thermometry," Proc. SPIE 8581, Photons

Plus Ultrasound: Imaging and Sensing 2013, 858118 (4 March 2013); doi:

10.1117/12.2004063

Event: SPIE BiOS, 2013, San Francisco, California, United States

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Single-cell photoacoustic thermometry

Liang Gao, Lidai Wang, Chiye Li, Yan Liu,

Haixin Ke, Chi Zhang, and Lihong V. Wang

Department of Biomedical Engineeri

ng, Washington University in St. Louis

One Brookings Dr., St. Louis, MO, 63130

ABSTRACT

A novel photoacoustic thermometric method is presented for

simultaneously imaging cells and sensing their temperature.

With 3 seconds per frame imaging speed, a temperature resolution of 0.2

o

C was achieved in a photo-thermal cell heating

experiment. Compared to other approaches, the photoacoustic thermometric method has the advantage of not requiring

custom-developed temperature-sensitive biosensors. This

feature should facilitate th

e conversion of single-cell

thermometry into a ro

utine lab tool and make it accessible to a mu

ch broader biological

research community.

Keywords:

Single-cell thermometry, temperature sensing, photoacoustic microscopy

1. INTRODUCTION

Cellular events – such as division, gene expression an

d enzyme reaction – are often accompanied by intracellular

temperature changes[1]. Accurately measuring intracellular te

mperature may provide new insights into cellular signaling

and metabolism. However, sensing intracellular temperature is

not a trivial task, especially at the single cell level.

Currently, there are two major approaches to intracellula

r temperature sensing: one uses micro- or nano-scale

thermocouples [2, 3]; the other employs temperature-sensitive fl

uorescent dyes [4], proteins [5], or nanoparticles [6].

Although the thermocouple-based approach features high temperature resolution (~0.1

o

C [3]), to detect intracellular

temperature changes, the tip of the ther

mocouple has to be inserted into cells

via a micromanipulation system. This

invasive operation may interrupt normal cell metabolic cycles and cause cell damage. In addition, only one cell can be

measured at a time. On the other hand,

the fluorescence-based approach reali

zes simultaneous imaging and temperature

sensing with ~0.5

o

C resolution [7]. However, most temperature-sens

itive fluorescent biosensors present problems, such

as sensitivity to solution pH values and potential toxicity to cells [8].

To overcome these limitations, we present a novel single-cell photoacoustic thermometric method for intracellular

temperature sensing. The system is based on high-resolution photoacoustic microscopy (PAM), which measures

ultrasound signals induced by light absorption. PAM has

achieved diffraction-limited optical resolution, and been

successfully applied in cellular imaging

applications, e.g., detecti

ng nanoparticle-targeted cancer cells [9] and label-free

imaging of cell nuclei [10]. In PAM, the acquired photoacoustic (PA) image amplitude is proportional to the initial

pressure rise p

0

at the absorber, induced by short-pulse laser excitation. The initial pressure rise is given by [11]

2

0

()

s

aa

p

pFF

C

βν

μ

==Γ

. (1)

Here,

β

is the thermal expansion coefficient,

v

s

is the speed of sound in the medium,

C

p

is the specific heat capacity at

constant pressure,

μ

a

is the optical absorption coefficient, and

F

is the optical fluence. The Gruneisen parameter

Γ

is

temperature dependent in water, as given by the empirical relation

A

BT

Γ

=+

, (2)

where

A

and

B

are constants, and

T

is the local temperature surrounding the

absorber in degrees Celsius. Substituting

Γ

from Eq. 2 into Eq. 1 gives

0

()

a

pABTF

μ

=+

. (3)

Consequently, by measuring the photoacoustic signal generated by the absorber, the local temperature can be detected.

Although photoacoustic thermometry has been employed in evaluating biological tissue temperatures [12], to our

knowledge, this is the first time it has been used for single cell temperature imaging.

Photons Plus Ultrasound: Imaging and Sensing 2013, edited by Alexander A. Oraevsky, Lihong V. Wang,

Proc. of SPIE Vol. 8581 858118-1

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TC

Calibration

UST

Lens

(a)

Heating pad

Temperature sensing

Heating cw laser

UST

Lens

AL

-

(b)

2. SYSTEM DESCRIPTION

A recently developed voice-coil PAM system [13] (Fig. 1)

was employed to acquire te

mperature-dependent cellular

images. The voice-coil PAM system consisted of a PA probe and

a raster scanner. In the PA probe, short laser pulses at

532 nm (10-200-532, Elforli

ght Ltd.) were fo

cused onto the sample surface thro

ugh a set of optics. The generated

photoacoustic signals were collected by an ultrasound lens (6 mm aperture, 0.5 NA in water) and received by a high-

frequency ultrasound transducer (V2022 BC, Olympus NDT). To

achieve high sensitivity, the optical and acoustic foci

were confocally and coaxially aligned. The photoacoustic

probe was mounted onto a voice-coil-based scanner ( VCS-

1010, Equipment Solutions) to implement fast imaging.

In experiments, the sample was scanned at a 10 Hz cross sectional imaging speed and a 1/3 Hz volumetric imaging

speed. The lateral resolution was measured as 3.4 μm. A higher spatial resolution is possible by increasing the optical

numerical aperture.

Fig. 1. System layout of PA-based cellular temperature sensing during (a) calibration and (b) photo-thermal heating. In

(b), the cell was heated by a 100-mW CW laser via a multim

ode fiber. The intracellular temperature was monitored by a

voice-coil PAM system in real time

. AL, acoustic lens; UST, ultrasoun

d transducer; TC, thermocouple.

2. RESULTS

To demonstrate single-cell photoacoustic thermometry, an in

tracellular temperature sensing

experiment was carried out

on HeLa cancer cells during photo-thermal heating. The HeLa

cells were loaded with iron oxide micro-particles (mean

diameter, 2.5 microns, PI21353, Fisher Scientific) as both a photoacoustic imaging contrast agent and a photo-thermal

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(a)

37 °C

1.6x10°,

1.55

(6

1.5

43 1.45

1.4

Q 1.35

1.3

1.25

(b)

300

250

fC

200

(1)

150 'A.

100

g

50

o

26 28 30 32 34

Temperature (Celcius)

(C)

heating source. The cells were immersed in PBS and imag

ed by the PAM. Laser pulse

energy was recorded by a

photodiode, and variations in the output were corrected pulse-by-pulse.

2.1 Sample preparison

The HeLa cells were maintained in Dulbecco's Modified Eagle Medium with 10% fetal bovine serum, 2 mM glutamine,

and 1% penicillin/streptomycin supplement. The cells were incubated at 37

o

C in 5% CO

2

, and were divided every 72

hours. After being dispersed in 0.25% EDTA-trypsin, they were seeded at 2-4×104 cells/cm

2

. As a PA imaging contrast

agent, iron oxide micro-particle

s (PI21353, Fisher Scientific; 1

μ

m size) were added to the cell medium in a

concentration of 0.5 pM and incubated for 24 hours. Coverslips with adherent cells were washed with PBS before

imaging.

Fig. 2. PA-based single-cell thermome

tric calibration. (a) Cell image at 23

o

C. (b) Cell image at 37

o

C. (c) PA amplitude

vs. temperature for the cell in (a).

The coefficient of determination,

R

2

, was 0.994.

2.2 Calibration

To study the relationship between intracellular temperature

and PA image amplitude, a coverslip with adherent HeLa

cells was uniformly heated by a custom-made heating pad, as shown in Fig. 1(a). The heating pad was made by

connecting a thin metallic platform to

a resistive heater (HT10K, Thorlabs). A

thermocouple was placed in contact with

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PA recovered temperature

(Celsius)

N

W

N

p

OC,J

Ui

4

6

....

A

_,

Ó

N

CD

3

CD

IV

O

_

....__Z

C,J '

O

.P

O

C.71

O

a)

O

0

N

CY1

the coverslip and close to (~ 1mm) the cells. The envir

onment temperature was gradually raised by increasing the

voltage applied to the heating pad. At each temperature, after

enough time had elapsed to ensure that thermal equilibrium

had been reached among the cells, the thermocouple, and the

environment, a PA image was

acquired and the temperature

was measured by the thermocouple. Since the thermal and stre

ss relaxation times for a 2.5 micron particle are 270 ns and

1.7 ns, respectively, which are longer than the laser pulse duration (~1 ns), both thermal and stress confinement

conditions were satisfied. After fluence correction, the acqu

ired PA image amplitude (peak-to-peak value of each A-line

signal) was given by

0

aa

p

Sk

kBTkA

F

μμ

== +

, (4)

where k is the system’s pressure sensitivity.

The measured PA signal vs. temperature for a single cel

l is shown in Fig. 2. The PA

signal of the cell was calculated

by averaging all pixels within the cellular boundary. Th

e result indicates that the cellular temperature rise was

accompanied by a linear increase

in its PA signal. The slope

a

kB

μ

and intercept

a

kA

μ

thus could be solved by linear

regression. Since the cellular temperatur

e has a one-to-one correspondence to its PA signal, the PA-recovered cellular

temperature was calculated as

a

IkA

T

kB

μ

−

=

. (5)

2.3 Single-cell temperature sens

ing during photo-thermal heating

Localized photo-thermal heating is widely used in cancer

therapy. Measurement of cellu

lar temperature during photo-

thermal heating is important to understanding the thermodynami

cs of cancer cells. To simula

te the heating process, the

output from a 100-mW CW laser (MLL-III-532, General Optoelectronic) was guided to the sample by a multimode fiber

(BFL37-600, Thorlabs) to heat the HeLa cancer cells (see Fig. 1(b)).

The experiment was divided into two steps. First, a control experiment without photo-thermal heating was performed to

provide a baseline. The cellular temperat

ure was monitored by the PAM at constant room temperature (Fig. 3). The

image acquisition speed was 3 seconds per frame. With this im

aging speed, the temperature resolution – defined as the

standard deviation of the measured control temperature – was 0.2

o

C. Note that by using a slower imaging speed or a

finer imaging step size, the temperature resolution can be furt

her improved. In the second step, the heating CW laser was

turned on and the cellular temperature was continuously measur

ed by the PAM at the same acquisition speed. As seen in

Fig. 3, the PA-recovered cellu

lar temperature rose from 23

o

C to 26

o

C during the 300 second heating period. After the

heating laser was turned off, the cellular temperatur

e gradually decreased to

the room temperature.

Fig. 3 Single cell temperature sensing during control and photo-thermal heating stages.

To test cells’ viability after these procedures, the cells we

re stained with green-fluoresc

ent calcein-AM (L-3224, Life

Technologies, Inc.) to indicate intracellular esterase activit

y and red-fluorescent ethidium

homodimer-1 to indicate loss

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of plasma membrane integrity. The cells were re-imaged on a fluorescence microscope (Fig. 4), and the results show that

> 80 % of cells were still viable after photoacoustic imaging.

Fig. 4. Cellular viability test after photoacoustic imaging.

Cells emitting green fluorescence are alive, while those

emitting red fluorescence are dead.

3. DISCUSSION

During raster scanning, repetitive probe laser pulses may

cause a local temperature rise. Here, this effect was

estimated by using the laser specific parameters and a ty

pical value of thermal diffusivity in water (1.4×10-7 m

2

·s-1 at 25

o

C). The calculation indicated that two adjacent laser pulse

s can raise the local temperature by as much as 0.1

o

C.

Summing ~1000 laser pulses scanning across a cell,

the estimated total temperature rise was ~0.25

o

C, which is much

less than the cell temperature. If a smalle

r temperature rise is desired, lower

laser pulse energy and a slower pulse

repetition rate may be used.

We also evaluated the dependence of the system’s pr

essure sensitivity on the coupling

water’s temperature. First,

since the ultrasound transducer was in

air and the room temperature was thermostatically controlled during the

experiments, the transducer’s pressure

sensitivity could be considered constant. Next, because the coupling water’s

temperature change might alter the speed of sound, acoustic fo

cal distance, and focal spot size, the system’s pressure

sensitivity could be affected. Here we estimated the acoustic focal distance and spot size at 25

o

C and 33

o

C, with an

assumption of uniform temperature distribution in the coupling water. The focal distance was calculated from

q

qw

v

lR

vv

=

−

, (8)

where

l

is the focal distance,

q

v

is the sound speed in acoustic lens (quarts),

w

v

is the sound speed in water, and

R

is the

radius of the acoustic lens. The focal diameter

d

is computed from

2

0.71

w

vl

d

f

D

=

,

(9)

where

f

is the acoustic frequency, and

D

is the outer diameter of the acoustic

lens. When temperature increased from 25

o

C to 33

o

C, the focal distance and the focal diameter decr

eased by 0.5% and 1.7%, respectively. By linear

approximation, the total temperature’s effect on the

system sensitivity was estimated as (1-0.005)×(1+0.017)

2

-1=2.9%.

At the same time, the measured PA amp

litude increased by 41%, which was 13 tim

es higher than the system’s pressure

sensitivity change. Therefore the effect of the coupling water’s

temperature change on the system’s pressure sensitivity

was negligible.

The single-cell photoacoustic thermometry requires that

the cell stays in the same solution and cellular environment

during calibration and temperature measurement experiments. For

in-vivo

studies, since the tissue surrounding the cell

may also contribute to the PA signal

[14], the correspondence between PA amp

litude and cellular temperature acquired

by

in-vitro

calibration experiments may not be applicable. To

reduce the effect of surrounding tissue on the

in-vivo

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cellular temperature measurement, one

can minimize the ratio of the tissue’s absorption to the sensing particle’s

absorption by choosing a specim

en-specific wavelength.

4. CONCLUSION

In summary, a novel PAM-based method is presented for single-cell temperature sensing. With 3 seconds/frame

imaging speed, the PAM-based method has achieved a temperature resolution of 0.2

o

C in a photo-thermal heating

experiment. To the best of our knowledge, this is the first

time that photoacoustic temperature sensing has been realized

at the single-cell level.

Compared to other cellular temperature sensing methods, the PAM-based approach has the advantage of not requiring

custom-developed temperature-sensitive biosensors. One can c

hoose commercially available absorptive dyes or particles

as the temperature sensitive agent in cellular imaging applica

tions. Although not demonstrated in this paper, by tuning

the imaging laser wavelength, endogenous cellular absorption contrasts, such as hemoglobin, melanin, lipid, DNA/RNA,

and protein, can also be employed for intracellular temperature sensing without cellular staining. The presented

photoacoustic thermometry should make

single-cell temperature sensing accessi

ble to a much broader biological

research community.

ACKNOWLEDGMENTS

This work was sponsored by the National Institutes of He

alth (NIH) under grants R01 EB000712, R01 EB008085, R01

CA134539, U54 CA136398, R01 CA157277, R01 CA159959 and DP1 EB016986. L. V. Wang has a financial interest

in Microphotoacoustics, Inc. and Endra, Inc.; however, neither provided support for this work.

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