1.

Introduction

Photoacoustic (PA) imaging combines the advantages of high contrast of light absorption, high resolution, and penetration depth of ultrasound, and has been investigated for breast cancer detection in recent years.^1,2 The performance characteristics of PA imaging systems include the maximum imaging depth, sensitivity, resolution, and contrast.^3,4 PA imaging has good sensitivity, resolution, and contrast, but the major limitation is the optical penetration depth.^3–5 Light intensity decreases exponentially with penetration depth because of strong scattering and absorption in biological tissues.^6,7 The optical window in the range of 700 to 900 nm allows relatively higher light penetration depth.^6,7 PA signals can be improved by contrast agents, such as silica-coated super paramagnetic iron oxide nanoparticles,⁸ system optimization, signal processing, or higher laser power.⁹ However, contrast agents have side effects, system optimization increases the complexity and cost, and higher laser power exposure induces potential damage to skin.

Manipulation of tissue temperature has been used for thermal ablation, hyperthermia, and imaging.^10–13 The normal body temperature of a healthy adult is $\sim 37°\mathrm{C}$.^14,15 Studies show that most normal healthy tissues such as liver, kidney, and muscles can withstand temperatures up to 44°C for 30 min.^10,16–18 The relationship between thermal exposure time and temperature in most normal healthy tissues has also been studied.^18,19 The reported results show that heating effects of 60-min exposure time at 43°C are equivalent to 30 (15) min of exposure at 44°C (45°C). These and other results indicate that there exist ranges of maximum safe exposure times at substantially higher temperatures than normal for most normal healthy tissues,^16,17,20 and thus the possibility of using high temperature for significant improvements in the PA signal strength and imaging depth. Thermal energy can be locally delivered by microwaves (MWs) in biological tissues.¹¹ An MW-energy applicator (antenna) is placed on the target site to increase the temperature surrounding the applicator inside the tissue. Coaxial slot antennas are the most popular tools for MW heating purposes due to their small dimensions and low cost.^21,22

Generally speaking, two kinds of laser sources are used to generate PA signals: (1) pulsed and (2) continuous wave (CW)-modulated laser sources. The pulsed laser is the conventional laser source, but CW lasers have the advantages of low cost, portability, and narrow band detection [i.e., high signal-to-noise ratio (SNR)].⁹ The photoacoustic radar (PAR) uses chirp signals for laser-beam modulation and matched filtering for signal processing.²³ Temperature can affect PAR signals, which become stronger at higher temperatures.^12,13 The imaging depth and SNR can be improved through uniform heating.^12,13 Most research reports of MW heating have been in applications of surgical ablation.^{21,22,24–27} Our earlier study¹² showed that “under uniform heating,” the thermally enhanced PA radar probe can increase the signal strength from ex vivo beef muscle by 13% and the imaging depth over 20% with a temperature increase from 37°C to 43°C. However, imaging improvement through uniform heating takes very long and can cause discomfort and inconvenience to patients. Thus, it is necessary to find a localized heating method for precise tissue targeting and shortening the exposure time. The present feasibility study investigates the thermal effects on PAR signal enhancement “under targeted (localized)” heating of biotissues in the same temperature range using MW heating. The tradeoff was found to be rapid (localized MW) heating at relatively low signal strength increase versus very slow uniform heating with significantly higher signal strength increase.

2.

Theoretical Background

PAR signals were generated by harmonically modulating light illumination on absorbers.²⁸ The response amplitude spectrum $\tilde{p}(\omega )$ of the generated acoustic pressure can be written in one dimension as^12,23,29

The spectrum of the LFM signal $r_t$ is

The Fresnel integrals $C(X)$ and $S(X)$ are defined by

It can be shown that

The amplitude of $\tilde{h}(\omega )$ and $\tilde{s}(\omega )$ can be derived as

Fresnel integrals can be approximated as³¹

For frequencies higher than 300 kHz, the Fresnel integrals are $\sim 1/2$. Therefore, the final matched filter output amplitude spectrum of two signals $\tilde{h}(\omega )$ and $\tilde{s}(\omega )$ is approximately³² given by

The output amplitude spectrum is proportional to the duration of the chirp and inversely proportional to the bandwidth of the chirp. For narrow frequency bandwidth $B$ and long duration time $K$ chirps, the amplitude increases.

Water plays an important role in biotissues. The total water content of human tissues can be up to 77%.³³ The Grüneisen parameter of water increases monotonically from 0.11 at 20°C to 0.24 at 45°C [shown in Fig. 1(a)]. The thermal expansion coefficient increases 107% from 20°C to 45°C.³⁴ The specific heat capacity of water is a mildly temperature-dependent parameter and the variation is less than 1% from 20°C to 45°C.³⁵ The speed of sound in water increases with temperature from 1482 to $1536\mathrm{m}/\mathrm{s}$, which is 3.6% increase from 20°C to 45°C.³⁶ Figure 1(b) shows a comparison of the temperature dependence of the parameters constituting the Grüneisen constant. It can be seen that the thermal expansion coefficient contributes much more than the other two parameters to the increase.

Fig. 1

(a) Grüneisen parameter of water dependence on temperature and (b) increase rate comparison of Grüneisen parameter elements of water from 20°C to 45°C.

Bovine muscles are used as samples for experimental validation. Myoglobin and hemoglobin are the main light absorbers in bovine muscle tissues and the concentration of myoglobin is much higher than hemoglobin:³⁷ Myoglobin content is $0.41\mathrm{mmol}/\mathrm{kg}$, whereas hemoglobin content is only $0.066\mathrm{mmol}/\mathrm{kg}$ in bovine muscle.³⁷ Myoglobin concentration can determine the color of muscle: higher myoglobin concentration yields darker color.³⁸ In ex vivo tissues, myoglobin and hemoglobin are present in oxygenated or deoxygenated forms, but deoxygenated concentrations are higher. On the surface of ex vivo tissues, the degree of oxygenated concentrations is higher. After myoglobin is oxygenated (oxymyoglobin), color changes from dark purple to bright red, with the muscle assuming a bright cherry-red color.³⁸

A typical composition of bovine muscle consists approximately of 75% water, 19% protein, 3% fat, 1% fiber, 1% carbohydrate, and 1% ash.^39,40 Temperature-dependent empirical thermal property models of individual biotissue-components have been introduced for the Grüneisen parameter estimation.^12,41 The Grüneisen parameter of a typical bovine muscle dependence on temperature has been studied from 20°C to 45°C.¹² In this temperature range, the increase is 67% for the Grüneisen parameter, 1.2% for the speed of sound, 0.33% for the specific heat capacity, and 64% for the thermal expansion coefficient.¹² The highest increase is, again, the thermal expansion coefficient.

The PAR signal changes with the Grüneisen parameter, and thus becomes temperature dependent.^42–44 Temperature rise in tissue depends on both heating power and time, and is also related to tissue properties and heat transfer to the vasculature.¹² The specific absorption rate (SAR) describes the time-averaged amount of MW power deposited per unit mass of tissue (W/kg). The SAR of electromagnetic energy incident on a tissue sample is defined by

The MW-generated heat is strongly dependent on the absorption coefficient of tissue, and on the frequency and intensity of the energy source.⁴⁶ Heat can be reduced by increasing perfusion with temperature.⁴⁷ Assuming that tissue is electromagnetically heterogeneous and acoustically homogeneous, the Pennes bioheat transfer equation was developed for roughly describing temperature elevation in biological tissues⁴⁷

For MW heating, the geometry is shown in Fig. 2(a), where the cylindrical coordinate system is used. $(r,z,t)$, $(r^{\prime },z^{\prime },\tau )$ are the observation and source coordinates, respectively. The translated temperature of the thermal field in the medium is written as

Fig. 2

(a) Experimental setup for MW heating and (b) geometry of MW antenna.

The initial and boundary conditions for MW heating are given by

For the geometry shown in Fig. 2(a), it is assumed that the absorbing medium is cylindrical and semi-infinite lengthwise, $0<r^{\prime }<b$ in the radial direction, and the penetration depth is $L=1/{\mu }^{\prime }$ in the axial direction, corresponding to the absorbing area of the MW energy. The GF is separated into the product of a radial and an axial function⁴⁹

The radial function $G_R$ satisfies

Tumors contain higher water content than normal healthy tissues,⁵¹ and thus they have higher electrical conductivity than low-water-content healthy tissue.^52–54 Studies have shown that their electrical conductivity is about 10% higher than normal tissue.^55,56 This suggests that MWs can raise the temperature of tumor tissues higher than that of the host mammary tissues, and thus generate higher contrast between tumors and normal tissues. The MW penetration depth is the inverse of the microwave absorption coefficient, where the electric field power is reduced by $1/{\mathrm{e}}^{246}$ at depths between 8.5 cm for bovine fat tissue and 1.1 cm for bovine muscle tissue at 2.45 GHz, respectively.^46,52,57 Penetration depths in other tissues have values between those of fat and muscle.

3.

Experimental Setups

The experimental setup for MW heating used in this work consisted of a MW generator (SSG-4000HP, Mini-circuit), a 40-dB power amplifier (ZHL-16W-43+, Mini-circuit), a coaxial slot antenna, thermocouples, and a data-acquisition device (USB-6221, NI) as shown in Fig. 2(a). The generated MW signals were amplified with a power amplifier and then sent to the coaxial slot antenna. Since water is a good absorber of MW energy, an external waveguide applicator has difficulty in delivering the MW energy to the absorber immersed in the water (our setup), so a single-slot coaxial antenna was employed for this research. This type of antenna can locally produce an electromagnetic field around the slots and heat the target area (deep inside the absorber) efficiently.^21,22,24 The electrical power supplied to the antenna was set at 2 W. The frequency was set at 2.45 GHz, the most popular frequency used for hyperthermia and heating applications.^21,22,24 The coaxial slot antenna was constructed from a RG402 semirigid coaxial copper cable as shown in Fig. 2(b). The outer conductor and core were both made of copper. A 2-mm ring-shaped slot was cut by removing a part of the outer conductor near the tip of the cable. The outer and inner conductors were connected at the tip by soldering. The reflected signal was sampled at the reflected port of a directional coupler, and measured by a network analyzer (E5063A, Agilent). The measured power loss was about $-12\mathrm{dB}$ (about 5% loss). Fine-wire thermocouples (K-type, Omega) were employed for temperature measurements. The measured data were collected by a USB DAQ device (USB-6221, NI), and then sent to a computer synchronized with the MW generator.

The experimental setup diagram for the MW-enhanced PAR system is shown in Fig. 3(a). A CW diode laser emitting at 800 nm was employed. The light beam size was 3.5 mm and was modulated by linear frequency waveform chirps (0.3 to 1.3 MHz, 1 ms in duration) generated by the signal-generation card NI PXI-5421 (National Instruments, Austin, Texas). A focused ultrasound transducer (Panametrics-NDT, V314) was utilized as a detector. The detected signals were amplified with a preamplifier (Panametrics-NDT, 5676) first, and then were sent to the digital data-acquisition card NI PXIe-5122 (National Instruments, Austin, Texas) in a computer. The sampling rate of the data-acquisition card NI PXIe-5122 was $100\mathrm{MS}/\mathrm{s}$, and the maximum input voltage was 10 V. Bovine muscle was purchased from a local store and was stored in a refrigerator at 4°C. Most experiments were performed within 3 days from the purchase. Fresh samples used for each experiment were cut from the same part. The sample dimensions were 30-mm long, 30-mm wide, and 20-mm thick. The samples were all wrapped in thin transparent plastic wrap during the experiments.

Fig. 3

(a) Experimental setup of the MW coupled PAR system and (b) measured temperature on a heated surface spot of bovine muscle versus time compared with the analytically estimated MW heating temperature.

4.

Results and Discussion

Figure 3(b) shows the comparison of the experimentally measured temperature in ex vivo bovine muscle of the MW-heated surface spots with the estimated temperature using Eq. (38). For bovine muscle at 37°C, the density is $1073.445\mathrm{kg}/{\mathrm{m}}^3$,^12,41 the specific heat capacity is $3632.7\mathrm{J}/(\mathrm{kg}·°\mathrm{C})$,^12,41 and the thermal conductivity is $0.5228\mathrm{W}/(\mathrm{m}·°\mathrm{C})$.^41,46 These parameters are all temperature dependent, but the variations are all less than 2% from 37°C to 45°C,^12,41 and are thus ignored in this study. The thermal diffusivity of bovine muscle can be calculated from the defining equation $\alpha =(k/\rho c_p)$. This yields $\alpha =1.34\times {10}^{-7}{\mathrm{m}}^2/\mathrm{s}$ in the range 37°C to 45°C, so variation with $T$ is ignored. For the 30-s heating period in Fig. 3(b), all the absorbed energy was assumed converted to heat ($\eta =100\%$), and the experimentally measured temperature at the heated surface spot was in excellent agreement with the analytically estimated temperature. The SAR was estimated to be $3.1\mathrm{W}/\mathrm{kg}$ (the best-fit value of the theory to the experimental results). The estimated error between measured and theoretical results was less than 10%. The temperature rose to 44.5°C from 37°C after 30-s of MW heating, a 20% increase. No visible damage was observed after the heating experiment.

Figure 4(a) shows the amplitude of the cross-correlation peak between the reference (by a uniform heating method) and the measured PAR signals under the influence of MW heating. The results showed that the PAR signal strength increased by about 8% with MW heating at ambient temperature 37°C and the SNR increased by 5%. The average elevated temperature during the PAR measurement was $\sim 42°\mathrm{C}$. In comparison, with uniform heating, the PAR-signal-amplitude increased up to 13% from 37°C to 42°C with the sample immersed in heated water. For MW heating-assisted PAR, the nonuniformly distributed temperature field in tissue was generated by the MW applicator, and at lower temperatures, the PA signals were weaker.¹² Therefore, the nonuniform temperature distribution and the lower heating efficiency under MW heating were the reasons for the less-efficient PAR signal increase. Theoretically, the SNR can be improved by a further 8% with a more efficient thermally assisted method,^12,13 the single most important criterion for such efficiency being achieving temperature uniformity. The experimental results of the imaging depth enhancement study are shown in Fig. 4(b). The sample was placed in a box filled with 0.47% intralipid solution simulating an optical scatterer as shown in Fig. 4(c).^12,58 The estimated scattering coefficient of the 0.47% intralipid solution was $0.014{\mathrm{mL}}^{-1}\mathrm{L}{\mathrm{mm}}^{-1}$ ($\text{concentration}\times \text{sample thickness}$) and the absorption coefficient was $0.003{\mathrm{mL}}^{-1}\mathrm{L}{\mathrm{mm}}^{-1}$.^12,58 The box was translated with microstages to attain different subsurface absorber depths. At 18 mm, the PAR signal was measured with large errors (20%), and the noise was high. With MW heating, the PAR signal became stronger (about 8% increase). The experimental results show that the imaging depth was increased by 11% with MW heating assistance.

Fig. 4

(a) Measured PAR cross-correlation amplitude peak from ex vivo bovine muscle with MW and uniform heating, (b) PAR signal imaging depth study under MW heating, and (c) experimental setup for PAR imaging depth study (the absorber surrounding medium was 0.47% intralipid solution).

Figure 5 shows an experimentally one-line-scanned image of a 4-mm wide ex vivo bovine muscle sample. Figure 5(a) is the picture of the sample as it was placed between clamps. Figure 5(b) shows a bright PAR horizontal-line image at ambient temperature 37°C with intralipid solution as the scatterer. The sample was placed 18 mm below the intralipid surface. The scanning signals were collected while moving the laser beam and transducer in tandem across the sample. Figure 5(c) shows the respective image generated with additional MW heating. Comparing Figs. 5(b) and 5(c), the MW-assisted image in Fig. 5(c) appears brighter than in Fig. 5(b) with enhanced contrast against a receding background due to the stronger signal (8% increase).

Fig. 5

(a) Picture of a 4-mm wide ex vivo bovine muscle sample, (b) one-dimensional image at 37°C with optical scattering, and (c) same imaging conditions as in (b) with added MW heating.

5.

Conclusions

A feasibility investigation of localized MW heating effects on PAR signal enhancement in ex vivo bovine muscle was conducted. The localized MW heating was experimentally and theoretically investigated. The calculated and measured results showed agreement at the heated spot over the same heating period and they demonstrated that MW sources can generate a thermal field at low-input-power in bovine muscle for enhancing PAR signals. The PAR SNR increased by about 5% with MW heating assistance for tissue temperature rise up to 42°C, whereas the maximum imaging depth was improved by 11%. MW heating was thus shown to be an effective heating source for biotissues. In view of earlier tissue heating studies,¹² it is concluded that MW heating can be used as a substantially more significant depth resolution enhancement method for PAR imaging of tissues, resulting in up to 13% increase in signal strength and over 20% increase in imaging depth under a localized temperature increase from 37°C to 43°C, provided the MW source can be optimized to generate more uniform subsurface tissue heating around the investigated tissue location than the system developed for this study.