3.1.

Experimental Setup

An integrated imaging and therapy system was assembled to acquire photoacoustic and ultrasound frames during photothermal therapy. The diagram of the experimental setup is presented in Fig. 1 . A 38-mm aperture, 128 element linear array transducer, with element spacing of about $300\mathrm{\mu }\mathrm{m}$ , (SonixRP, Ultrasonix Medical Corporation, Canada) was used to capture ultrasound pulse-echo and photoacoustic transients. The signals were recorded using a $40\text{-}\mathrm{MHz}$ sampling rate. The nominal center frequency of the transducer in the pulse-echo regime was $5\mathrm{MHz}$ . An OPO pulsed laser system (Vibrant B, Opotek Incorporated, USA) operating at an $800\text{-}\mathrm{nm}$ wavelength, with a $5\text{-}\mathrm{ns}$ pulse duration and a $10\text{-}\mathrm{Hz}$ repetition rate—providing optical fluence up to $15\mathrm{mJ}∕{\mathrm{cm}}^2$ —was interfaced for photoacoustic imaging of tissue samples. An Nd:YAG pulsed laser (Polaris II, New Wave Research, Incorporated, USA) operating at a $532\text{-}\mathrm{nm}$ wavelength, with a $5\text{-}\mathrm{ns}$ pulse duration and a $20\text{-}\mathrm{Hz}$ repetition rate, was used in the tissue-mimicking phantom experiments. The direction of the laser beam was orthogonal to the imaging plane, as shown in Fig. 1.

Fig. 1

Experimental setup for photoacoustic and ultrasound imaging during therapy.

During imaging, the sample was first irradiated using a pulsed laser, and the photoacoustic response was received on one element of the transducer array. The laser pulses were repeated until signals from each transducer element were collected. Immediately after the photoacoustic imaging, conventional pulse-echo ultrasound imaging was performed. The received photoacoustic and ultrasound signals were then used to reconstruct corresponding images using a delay-and-sum beamforming approach.³⁵ Since both ultrasound and photoacoustic data were acquired by the same ultrasound transducer array, the reconstructed images were spatially coregistered.

A continuous wave diode laser (HAM, Power Technology, Incorporated, USA), operating at $800\mathrm{nm}$ with a maximum power of $1\mathrm{W}$ , was used as a light source for the photothermal therapy. During the four-minute exposure, ultrasound and photoacoustic frames were recorded every $10\mathrm{s}$ . The captured data were stored offline for temperature processing. The experiments were performed at a room temperature of $25°\mathrm{C}$ .

3.2.

Synthesis of Photoabsorbers

Gold nanoparticles of $20-\mathrm{nm}\mathrm{diam}$ were prepared from chloroauric acid by using sodium citrate as a reducing agent.³⁶ The extinction peak of the particles, determined by UV-Vis spectroscopy, was around $532\mathrm{nm}$ . For ex-vivo tissue studies, $90\text{-}\mathrm{nm}\mathrm{diam}$ dextran–stabilized composite nanoparticles were designed to have a broad absorption, between 500 to $1000\mathrm{nm}$ wavelength.³⁷ To prepare the composite nanoparticles, iron oxide particles were first synthesized by a coprecipitation of ${\mathrm{Fe}}^{2+}$ and ${\mathrm{Fe}}^{3+}$ ions in alkali conditions.³⁸ Gold shells were then coated on the iron oxide nanoparticles by reducing the chloroauric acid via iterative seeding and reduction methods.^{39, 40} The aqueous composite nanoparticles were stabilized by dextran at the pH value of 5.0.

3.3.

Sample Preparation

Two tissue-mimicking phantoms, measuring $50\times 50\times 50{\mathrm{mm}}^3$ $(\mathrm{W}\times \mathrm{L}\times \mathrm{H})$ , were prepared using polyvinyl alcohol (PVA). PVA has modest optical absorption, scatters light similar to tissues, and has been used in constructing phantoms for optical imaging studies.⁴¹ Additionally, PVA and tissue have a similar speed of sound.⁴¹ To fabricate phantoms, an 8% aqueous PVA solution was poured into a mold and set to the desired shape by applying two freeze and thaw cycles.⁴² Acoustic contrast was obtained by adding 1.0% (by weight) silica particles of $40\mathrm{\mu }\mathrm{m}$ diam to the PVA solution. Gold nanoparticles ( $0.096\text{-}\mathrm{mg}\mathrm{Au}∕\mathrm{mL}$ solution) were embedded in one phantom, while the second phantom did not have any photoabsorbers added.

Ex-vivo tissue imaging and temperature monitoring studies were performed using a fresh porcine muscle specimen. The $30\times 30\times 15\text{-}{\mathrm{mm}}^3$ samples were immersed in water for acoustic coupling between the ultrasound probe and the tissue. $20\mathrm{\mu }\mathrm{L}$ of composite nanoparticles ( $0.057\text{-}\mathrm{mg}\mathrm{Au}∕\mathrm{mL}$ solution) were injected using a 23-gauge hypodermic needle, at depths of 8 to $16\mathrm{mm}$ from the tissue surface, under ultrasound guidance. The needle was inserted so that it was orthogonal to both the ultrasound imaging plane and the laser beam, thereby minimizing interference with the measurements. The injection lasted about $12\mathrm{s}$ long, after which imaging and/or therapy began immediately to prevent passive diffusion of the photoabsorbers in the tissue.

3.4.

Temperature Monitoring

Prior to the remote measurements of the temperature, tissue phantoms and animal tissue samples were calibrated to establish the relationships between temperature and the relative change in photoacoustic signal amplitude or temperature, and the relative time shift of ultrasound signals. The samples were immersed in a water tank capable of maintaining a user-defined constant temperature. A thermistor was inserted into the center of the sample to directly and independently measure the temperature. First, a baseline ultrasound and photoacoustic frames were acquired. The water tank temperature was then increased in $1°\mathrm{C}$ increments from 25 to $35°\mathrm{C}$ . The system was allowed to reach equilibrium at each temperature, and photoacoustic and ultrasound frames were recorded.

The increase in the photoacoustic signal amplitude was calculated between successive photoacoustic frames. Before comparing the successive photoacoustic signals, the apparent time shift in the photoacoustic signal due to thermally induced speed of sound change was accounted for. A normalized change in the photoacoustic signal versus the temperature relationship was obtained for the sample. During therapy, the change in the signal amplitude was calculated between successive photoacoustic frames, and the temperature rise was computed from the measured photoacoustic signal amplitude-temperature relationship.

Apparent time shifts between successive ultrasound frames were calculated using a cross-correlation-based, motion tracking algorithm.⁴³ Axial strain was then estimated by differentiating the time shifts along the temporal (i.e., axial) direction. Thus, a strain versus temperature dependence was obtained for the sample. While performing the photothermal therapy procedure, the axial strain between successive ultrasound frames was estimated, and the previously measured strain-temperature relationship was used to compute temperature.^{10, 44}

Signal-to-noise ratio (SNR) and contrast-to-noise ratio (CNR) were used to quantify the quality of the photoacoustic and ultrasound thermal images. The SNR was calculated using the following expression:

4.1.

Detection of Photoabsorbers

The ultrasound and photoacoustic images in Fig. 2 represent a $20\times 20\text{-}\mathrm{mm}$ field of view of a porcine muscle tissue before and after the injection of composite nanoparticles.

Fig. 2

(a) Ultrasound image of tissue sample injected with photoabsorbers. Photoacoustic images (b) before and (c) after the injection shows the enhanced optical contrast. (d) Combined ultrasound and photoacoustic image after injection. All images cover a $20\times 20\text{-}\mathrm{mm}$ field of view.

The injected photoabsorbers are not visible in the ultrasound image [Fig. 2a]. Photoacoustic images before [Fig. 2b] and after [Fig. 2c] the injection show the photoacoustic signal change due to the local injection of the photoabsorbers. The overlaid ultrasound and photoacoustic image demonstrates the complementary characteristic of each imaging technique – the structure of the tissue displayed in the ultrasound image and differences in the optical contrast, artificially enhanced by photoabsorbers, illustrated in the photoacoustic image. Additionally, since photoacoustic imaging was performed at $800\mathrm{nm}$ , i.e., in the NIR spectrum, the signal from the porcine tissue is relatively small. Photoacoustic transients from the tissue at or near the injection site [area between arrows in Fig. 2d] before and after the photoabsorber injection, are compared in Fig. 3 . Overall, the two signals are not very different, except at the injection site, where there is a noticeable increase of the photoacoustic signal.

Fig. 3

Photoacoustic RF signals from the injection site between the arrows in Fig. 2d.

4.2.

Temperature Calibration

Images in Fig. 4 show the ultrasound and photoacoustic images of the porcine tissue at the initial temperature of $25°\mathrm{C}$ . During the measurements, the tissue sample was completely immersed in the water tank. Beamformed photoacoustic signals collected at different time/temperatures from the region, denoted by the arrows in Fig. 4b, are plotted in Fig. 4c. There is a steady monotonic increase in the photoacoustic signal intensity as the temperature increases. Overall, a 42% increase in photoacoustic signal intensity (Fig. 5 ) was observed for a $9°\mathrm{C}$ change in temperature.

Fig. 4

(a) Ultrasound and (b) photoacoustic image of a porcine muscle tissue covering a $15\times 15\text{-}\mathrm{mm}$ field of view. (c) Successive beamformed photoacoustic A-lines plotted for increasing temperature demonstrated both amplitude increase and temporal shift. Note that (b) the region between the arrows is plotted covering a $10\text{-}\mathrm{mm}$ axial distance.

Fig. 5

Comparison of photoacoustic signal increase and ultrasound strain change for the same temperature rise. The error bars represent the standard deviations obtained from ten measurements.

Along with amplitude change, the photoacoustic signal [Fig. 4c] shifts and appears closer to the transducer as the temperature increases. This time shift is due to the thermally induced speed of sound change. The change in the speed of sound with temperature was used in the ultrasound-based thermal imaging, where the apparent time shifts of the ultrasound pulse-echo signals were measured to assess temperature. Compared to the photoacoustic signal change, the measured strain (Fig. 5) in the ultrasound calibration was less than 1% for the same temperature rise. Therefore, photoacoustic thermal imaging has a greater change in the measured signal when compared to the ultrasound imaging. Furthermore, the SNR was calculated at each incremental temperature for both thermal images. The average SNR for the ultrasound-based strain images was $45.56\mathrm{dB}$ , while the average for normalized photoacoustic images was $59.97\mathrm{dB}$ . The average standard deviation for the ultrasound signal change was 0.003%, which corresponds to a temperature resolution of $0.05°\mathrm{C}$ . The average standard deviation for the photoacoustic signal change was 1%, which led to a temperature resolution of $0.16°\mathrm{C}$ . However, temperature resolution of photoacoustic thermal imaging can be improved by monitoring the energy of each laser pulse. Indeed, the temperature resolution of the photoacoustic imaging was largely determined by the pulse-to-pulse stability of the laser source. Furthermore, temperature resolution can be increased by averaging several laser pulses.

To examine the effect of the photoabsorbers on the temperature dependence of the photoacoustic amplitude, the temperature calibration procedure was performed on two similar PVA phantoms. The first phantom did not contain any nanoparticles and was used as a control, while the second phantom had gold nanoparticles added. Photoacoustic imaging was performed at $532\mathrm{nm}$ to match the absorption resonance of the photoabsorbers. The normalized change in the photoacoustic signal (Fig. 6 ), for a $9°\mathrm{C}$ increase in temperature, was nearly identical in both phantoms, demonstrating that gold nanoparticles have no effect on the slope of the measured change in signal. However, the addition of photoabsorbers increases the measured photoacoustic signal (Fig. 3), as observed in the tissue experiments. Thus, as expected, the average SNR in the nanoparticle-embedded phantom, $72.58\mathrm{dB}$ , was higher than the average SNR in the plain phantom, $49.01\mathrm{dB}$ .

Fig. 6

Photoacoustic signal change for a $9°\mathrm{C}$ temperature increase in phantoms with and without photoabsorbing nanoparticles.

4.3.

Therapy Guidance

Experimental photothermal therapy was carried out for four minutes on fresh porcine muscle tissue to evaluate ultrasound and photoacoustic thermal imaging. The initial ultrasound and photoacoustic images are displayed in Figs. 7a and 7b , covering a $20\times 20\text{-}\mathrm{mm}$ region. The site of the photoabsorber injection is indicated by the dashed circle in the ultrasound image, which also corresponds to the location of the strong photoacoustic response.

Fig. 7

Photothermal therapy experiment performed on porcine tissue with injected photoabsorbers. (a) Initial ultrasound and (b) photoacoustic images. The injection site is marked by a dashed circle in the ultrasound image. (c) through (f) Photoacoustic-based thermal images and (g) through (j) ultrasound-based thermal images after 60, 120, 180, and $240\mathrm{s}$ of therapy. All images cover a $20\times 20\text{-}\mathrm{mm}$ field of view.

The normalized photoacoustic and strain images obtained during the therapeutic procedure were converted to temperature maps using the relationship between the temperature and either the photoacoustic pressure, or the relative time shift (Fig. 5), respectively. Before estimating temperature, the time shift that occurs with the photoacoustic signals, due to speed of sound, was compensated for. The photoacoustic thermal images [Figs. 7c, 7d, 7e, 7f] and ultrasound thermal images [Figs. 7g, 7h, 7i, 7j], which were computed after 60, 120, 180, and $240\mathrm{s}$ of therapeutic laser radiation, show the progressive increase in temperature. After four minutes of therapy, the region with injected photoabsorbers reached a temperature elevation of slightly less than $10°\mathrm{C}$ , while the surrounding region had a temperature rise of less than $2°\mathrm{C}$ . The therapeutic region is spatially and temporally coregistered in both ultrasound and photoacoustic thermal images.

The temperature elevation obtained by photoacoustic and ultrasound thermal imaging was compared in a $1\times 1\text{-}\mathrm{mm}$ region, approximately centered in the therapeutic zone. The temperature rise computed by both methods was highly correlated [Fig. 8a ] throughout the procedure. The maximum temperature difference between the two imaging techniques was less than $0.5°\mathrm{C}$ , while the mean absolute temperature difference over the four minute procedure was $0.26°\mathrm{C}$ . However, the photoacoustic thermal imaging had a higher CNR compared to the ultrasound imaging [Fig. 8b]. The CNR for both techniques increased as the temperature increased. Indeed, a higher temperature in the therapy region leads to a greater photoacoustic signal as well as higher strain. Overall, the results indicate temperature monitoring during an experimental photothermal therapy procedure can be performed by both photoacoustic and ultrasound thermal imaging simultaneously.

Fig. 8

(a) Comparison of photoacoustic and ultrasound temperature measurements in a $1\times 1\text{-}\mathrm{mm}$ region inside the therapy zone. (b) Contrast-to-noise ratio (CNR) measurements indicate that the CNR increases with temperature for both thermal imaging techniques.