2.1.

Shaping of the Detection Laser Pulses

Using laser beams to generate and detect ultrasound in biological tissues immediately leads to opposite requirements. First, the sensitivity of the measurement can only be increased by using a higher power of laser radiation. Second, the laser power is limited by safety limits; namely, the maximum permissible exposure (MPE), to avoid damaging the biological tissue.²⁰ Those opposite requirements necessitate an appropriate laser exposure management.

Efficient laser generation of acoustic waves is obtained with sufficiently short laser pulses compared with thermal and stress relaxation times.³ Typically, few nanosecond pulses produced by Q-switched lasers are appropriate. In practice, the tissue is exposed repetitively over a large surface area, and the average intensity is limited by the MPE for a repetitive laser exposure.²⁰

The power management is mostly related to the detection laser. When probing soft tissues at depths limited to 1 or 2 cm, propagation delays of ultrasonic waves are below 20 μs since the propagation velocity is about $1.5\mathrm{mm}/\mu \mathrm{s}$. The detection laser illumination should thus be limited to the propagation delay to minimize laser exposure. For most industrial applications, a master oscillator (MO) power amplifier emitting powerful long pulses is appropriate. Such a laser source can be adapted to biomedical applications by limiting the pulse duration without losing energy in the laser system. This is done by adding an intensity modulator (IMOD) between the continuous wave (cw) MO and the power amplifier. A schematic diagram of such a laser source is shown in Fig. 1(a). The IMOD allows inhibiting the amplifier output at the beginning of each pumping flash by setting its transmission to zero. The energy then builds up in the amplifier laser rods (LRs). Before the end of each pumping flash, the IMOD transmission is increased gradually to release the stored energy and obtain the desired pulse duration and shape at the amplifier output. Without pulse shaping [dashed line in Fig. 1(b)], the laser pulse has an energy of 35 mJ, a peak power of 280 W, and a duration of 110 μs (FWHM: full width at half maximum). With pulse shaping [solid line in Fig. 1(b)], the laser pulse tailored for this experiment has an energy of 14 mJ, a peak power of 560 W, and a duration of 24 μs (FWHM). Reducing the pulse duration thus increases the available peak power. More importantly, the peak power for a given energy is about five times higher with the tailored pulse. The tailored pulse was shaped with an increasing instantaneous power for two reasons. First, this eliminates strong signal oscillations at the output of high-pass filters preceding signal digitization. Second, the sensitivity increases with time to facilitate the detection of weaker signals arriving at larger time delays (coming from larger depths in the tissue).

Fig. 1

Detection laser. (a) Layout of the system. MO, Nd:YAG cw master oscillator; ISO, optical isolator; HW, half-wave plate; PBS, polarizing beam splitter; AM, amplitude modulator; IMOD, intensity modulator; TFP, thin-film polarizer; LR, Nd:YAG laser rod; FL, flashlamp; QW, quarter-wave plate. Other components are plane dielectric mirrors. (b) Natural pulse shape of the detection laser (dashed line) and tailored pulse shape (solid line) used in the experiment. The generation laser pulse occurs at $t=0\mu \mathrm{s}$.

2.2.

Scanning and Detection Systems

A schematic diagram of the scanning setup is shown in Fig. 2(a). The generation laser (not shown) is a Q-switched laser emitting 5 ns pulses at 532 nm with a repetition frequency of 10 Hz. The generation laser beam is directed toward the biological tissue specimen after beam expansion and transmission through a line mask, blocking the generation laser on the detection scan line. The detection laser beam is transmitted to the setup by a multimode optical fiber. A sample of the detection laser beam is first deflected by a beam splitter (BS) and coupled into an optical fiber as a reference beam for phase-noise elimination. The detection laser beam is then focused on a 400-μm-diameter spot on the tissue specimen. A small prism mirror ($P$) in front of the collecting optics is used to direct the detection beam to the tissue without coupling stray light into the collecting optics. The resulting illumination/detection geometry is shown in the inset [Fig. 2(a)] where the elliptic area is the generation laser spot (masked on the scan line), and the centered circular area is the detection laser spot. The line mask eliminates any overlap between illumination and detection, thus allowing an independent evaluation of the MPE for generation and detection lasers. The collecting optical system uses two lens assemblies (LAs) for collimating and coupling the reflected/backscattered light into the signal beam optical fiber (400-µm-core diameter, 0.39 numerical aperture). The biological tissue specimen is located on a computer-controlled motorized stage, allowing displacements along the scanning axis $x$ and the focusing axis $z$.

Fig. 2

Experimental setup. (a) Schematic diagram of the biological tissue scanning setup. LA, lens assembly; L, lens; M, plane mirror; BS, beam splitter; P, gold coated hypotenuse prism. (b) Schematic diagram of the CFPI in differential configuration. PBS, polarizing beam splitter; PZM, piezoelectrically actuated mirror mount; VA, variable attenuator; DD, differential detector.

The phase demodulation of the collected light is done with an actively stabilized CFPI mounted in differential configuration to achieve both intensity and phase-noise reduction [Fig. 2(b)].²⁴ The horizontal polarization is used to demodulate light from the tissue specimen (signal beam), and the vertical polarization is used to measure the phase noise of the detection laser (reference beam). A CFPI has been preferred to a PRI since a PRI needs a longer illumination of the biological tissue to write the adaptive volume grating inside the photorefractive crystal. This longer illumination implies that more energy is delivered to the tissue for a given peak power. Although this is not critical for material,²¹ the shortest illumination time is preferable in biomedical applications where the exposure is limited by safety limits.²⁰ The CFPI is a 1-m-long cavity with mirrors’ reflectivity of 94.5%. The back mirror is mounted on a piezoelectrically actuated mount (PZM) fed by a servo loop ensuring stabilization at half-maximum of the carrier frequency transmission. This provides a peak demodulation response at about 2 MHz with an acceptable response (larger than 30% of the peak) between 400 kHz and 9 MHz.^17,22 For both reference and signal beams, a differential detector (DD) is used for intensity noise reduction. Each DD uses two InGaAs photodiodes in series DC-coupled to a transimpedance amplifier followed by a high-pass filter (300-kHz cutoff frequency) and a voltage amplifier. Reference and signal channels are digitized and processed to eliminate the phase noise of the laser source by the adaptive filtering explained as follows. First, the amplitude of the signal channel before the generation laser pulse was multiplied to obtain an rms noise equal to that of the reference channel in the same time window. Then iterations around this first value have allowed determining the optimum amplitude multiplication factor to minimize the residual noise on the difference between both channels. Typically, the residual noise was comparable with the shot noise level calculated from nominal values of the transimpedance gain and InGaAs photodiodes’ quantum efficiency. The spectrum of the residual noise was practically white in the frequency range kept for the analysis (0.5 to 3 MHz). The adaptive filter also included a fixed numerical delay between both channels to account for any electronic delay in the setup.

2.3.

Image Reconstruction

A typical image is obtained from 60 to 80 A-scans, each measured with a single detection laser pulse (no averaging). All A-scans are grouped in a raw B-scan image, which is processed with the synthetic aperture focusing technique (SAFT).^3,25 The SAFT algorithm is applied in the time domain and takes into account the surface profile of the biological tissue. When measuring remotely on curved surfaces with non-negligible height variations compared with the minimum ultrasonic wavelength (about 0.5 mm at 3 MHz), the surface profile is an essential input for a reliable image reconstruction. In the experiment described here, unlike in the one previously reported, in which the $z$-scan approach was used,²² the surface profile was measured prior to the scan with an OCT system. Each raw B-scan image, once linked with the surface profile, contains all the necessary information for both NCPAT and NCUS imaging modes. This is due to two concomitant ways of generating ultrasound in the specimen. When the generation laser pulse illuminates the specimen, a significant proportion of the incoming light is diffusely reflected. This light penetrates only at shallow depths and is lost in free space after a few scattering events. The remaining light penetrates more deeply in the specimen by following highly randomized optical paths. The fluence associated to these penetrating photons decreases rapidly with depth.³

Light penetrating more deeply into the specimen contributes essentially to the photoacoustic generation on optical absorbers (e.g., blood vessels), the strength of the thermoelastic generation occurring on each absorber being related to the product of the local laser fluence, and its absorption coefficient at the generation laser wavelength (GLW). Ultrasonic waves generated in this process then propagate up to the surface of the tissue where they can be detected. The only difference with conventional PAT is the non-contact optical detection and the fact that the surface profile must be taken into account in the reconstruction, which is not generally needed in conventional PAT due to the known position of transducer($s$) and the close acoustic impedances of water and tissue.

The high laser fluence present at shallow depths contributes to the thermoelastic generation of an ultrasonic wave at the surface (or close to it) by the background absorption of the tissue. Even though the background absorption is usually low (when the GLW is properly chosen), the product of the high laser fluence at low depth with the low background absorption of the tissue is usually sufficient to generate an acoustic waves in a similar manner to industrial LU where the penetration depth of photons is usually much shorter. The ultrasonic wave generated by this process has an initial area equal to that of the generation laser spot and a wavefront following essentially the surface topography of the specimen over this area. This ultrasonic wave then propagates inside the tissue, and any acoustic impedance mismatch will reflect or backscatter ultrasonic waves toward the surface were they can be detected.

Consequently, the same data can be used for both NCPAT and NCUS, the only difference being the propagation delay, as schematically shown in Fig. 3. In the case of NCPAT [Fig. 3(a)], the delay is the propagation time from the absorbing inclusion to the detection location at the surface. In the case of NCUS [Fig. 3(b)], the delay includes, in addition, the propagation time from the surface to the acoustically mismatched inclusion.

Fig. 3

Reconstruction methods. (a) One-way path used in NCPAT imaging mode. (b) Two-way path used in NCUS imaging mode. G, generation laser beam; D, detection laser beam; S, signal beam.

The relative strength of both generation processes depends on the tissue optical properties and the chosen GLW. If the background absorption of the tissue is strong at the GLW, the ultrasonic wave generated at the surface will be strong, and less light will be available for photoacoustic generation at absorbing sites inside the tissue. On the contrary, if the background absorption of the tissue is very low at the GLW, the thermoelastic generation at the surface will be very low and more light will be available inside the specimen for thermoelastic generation on optical absorbers. In general, both generation processes will occur concomitantly. Of course, if the choice of the GLW strongly favors one mode of generation, this will be detrimental to the other. Most of the time, however, the technique is intrinsically bimodal; the same data being used for both processed images. The only difference is the expression for the temporal delay: the one-way propagation delay is used for NCPAT reconstruction, and the two-way propagation delay is used for NCUS reconstruction. When used as a bimodal technique, the photoacoustic signals may lead to artifacts in the NCUS image and the ultrasonic signals may lead to artifacts in the NCPAT image. However, these artifacts will generally be faint since SAFT reconstruction will not properly focus signal of one mode into the other one. More advance processing may also be used to minimize these artifacts, which can be seen as a kind of crosstalk.

The algorithm used for image reconstruction was time-domain SAFT. After applying adaptive filtering for noise reduction, each A-scan was numerically filtered with a third-order Bessel-type bandpass filter using typical cutoff frequencies of 0.5 and 3 MHz. With a lateral step size of 400 μm, using frequencies higher than 3 MHz only increases the background noise without a significant gain in spatial resolution.²⁶ Each A-scan was delayed according to the measured surface profile and to the speed of sound in the specimen (typically $1.55\mathrm{mm}/\mu \mathrm{s}$). The reconstruction grid included interpolation along the axis $x$ to reduce the lateral pixel size to 100 μm instead of the experimental lateral step of 400 μm. The vertical size of the pixel was set to 100 μm in the reconstruction. Consequently, processed images were obtained with a pixel size of $100\times 100{\mu \mathrm{m}}^2$. For both imaging mode, the time derivative of each A-scan was used to account for the fact that pressure is proportional to the velocity of the surface, not to its displacement.²¹ Although this is an approximate procedure, image quality was found to be improved in this way. Taking the true frequency response of a CFPI would be more accurate.^17,22

In the present work, the OCT measurements were performed prior to the scan but the working distance of the OCT system was similar to that of our setup (25 mm). Consequently, both systems could be integrated using a dichroic mirror. In this first implementation, a B-scan of 80 A-scans can be measured in 8 s. The processing time for NCPAT and NCUS was below 1 s. Consequently, measurement and processing time below 10 s is achievable without exceeding MPE for results shown in the following section.

2.4.

Preparation of Tissues and Exogenous Inclusions

Chicken breasts were cut parallel to the natural surface to obtain a uniform thickness of about 10 mm for the upper part of the breast. This procedure was used to ensure that the natural surface of the upper piece was intact (without any incisions or preparation) for laser measurements. In the lower piece of the breast, 1- to 2-mm-deep incisions were made with a scalpel and gently filled with vegetable oil using a syringe before inserting exogenous inclusions described below. Vegetable oil was also poured between both pieces of chicken breast for a better ultrasonic contact when the upper piece was put back over the lower piece. Calf brains were prepared similarly, thus keeping intact the natural upper surface of the brain for laser probing.

Blood vessels were simulated with polyester thin wall tubes filled with India ink diluted to obtain an absorption coefficient ${\mathrm{\mu }}_a$ comparable with that of whole blood at 532 nm, which is about $235{\mathrm{cm}}^{-1}$ for oxyhemoglobin and $217{\mathrm{cm}}^{-1}$ for deoxyhemoglobin. Those values were obtained by assuming a whole blood concentration of hemoglobin of $150g\mathrm{Hb}/l$.²⁷ The polyester tubes were optically transparent at visible and near-infrared wavelengths and the nominal wall thickness was 12.7 μm. The negligible wall thickness compared with the shortest ultrasound wavelength considered in the experiment (about 0.5 mm) also ensured a good acoustic transparency. Consequently, these phantoms were appropriate to mimic the optical absorption of blood vessels while minimizing the ultrasonic impedance mismatch with the surrounding tissue.

Grayish metal wires were also used as exogenous inclusions to provide a strong ultrasonic impedance mismatch with the surrounding tissue. When used without white paint, their grayish color also ensured some optical absorption (to generate a photoacoustic signal). When painted white, their optical absorption was essentially eliminated, and no photoacoustic signal was expected in this case.

2.5.

Laser Safety Limits and Sensitivity

For the generation laser, the average intensity must be considered since there is a significant overlap of the surface illuminated by the generation laser between successive points of the scan. The illuminated elliptic surface of $4.0\mathrm{cm}\times 2.5\mathrm{cm}$ has an area of $7.9{\mathrm{cm}}^2$. With 100 mJ pulses at 10 Hz, this corresponds to an average intensity of $130\mathrm{mW}/{\mathrm{cm}}^2$, well below the MPE at 532 nm ($200\mathrm{mW}/{\mathrm{cm}}^2$).²⁰ The single-shot MPE, expressed in terms of fluence, can be used at the detection laser wavelength (1064 nm) since there is no overlap between successive points of the scan. Assuming a pulse duration $t_p$ of about 25 μs, the MPE is given by $C_A$ $t_p^{0.25}=0.39\mathrm{J}/{\mathrm{cm}}^2$ since $C_A=5.0\mathrm{J}/{\mathrm{cm}}^2/{\mathrm{s}}^{0.25}$ at 1064 nm.²⁰ The 400-μm-diameter detection spot corresponds with a surface $S=1.3\times {10}^{-3}{\mathrm{cm}}^2$. The maximum energy per pulse is thus $E_{\max }=\mathrm{MPE}\times S=0.5\mathrm{mJ}$, which was the pulse energy used in the experiment.

The sensitivity can be estimated from the shot-noise limit, which was essentially reached after adaptive filtering. Considering the amount of light captured by detectors, we have estimated a noise equivalent pressure of 8 Pa for a bandwidth of 3 MHz (corresponding with a surface displacement of about 0.8 pm).²² Some passive losses could be reduced further in our setup, but this first estimate gives the right order of magnitude. From the derivation performed in Ref. 28, we calculate a noise equivalent pressure of about 6 Pa for a 0.4-mm-diameter polyvinylidene fluoride (PVDF) transducer. Therefore, the sensitivity obtained during the experiment was comparable with the theoretical sensitivity of a piezoelectric transducer.*

3.1.

Images of a Chicken Breast Specimen

Results obtained with a chicken breast specimen are shown in Fig. 4. As mentioned in Sec. 2.4, embedded objects were chosen to produce different combinations of photoacoustic and ultrasonic responses. The NCPAT image [Fig. 4(a)] and the NCUS image [Fig. 4(b)] of the same chicken breast specimen clearly exhibits the embedded objects zoomed below their respective image. The 0.7-mm-diameter grayish metal wire (i) is seen in both NCPAT and NCUS images. Blood vessel phantoms with diameters of 0.5 mm (ii), 0.3 mm (iii), and 0.7 mm (vi) are mainly seen in the NCPAT image. The 0.8-mm-diameter white-painted metal wire (v) is only seen in the NCUS image. Many additional dark spots seen in the NCPAT image are attributed to endogenous absorption sites. These darker spots should not be confused with noise spikes since successive measurements on the same chicken breast specimen gave reproducible results for both exogenous and endogenous absorption sites. Faint hyperbolic artifacts seen in Fig. 4(b), especially at shallow depths, are attributed to photoacoustic signals which are not properly focused by the NCUS reconstruction algorithm. The lateral amplitude profiles (along $x$) extracted for each embedded objects are also shown in Fig. 4 using the same vertical scale to exhibit both the lateral resolution and the relative signal strength. For the NCUS reconstruction, a proper modeling of the source near the surface and the following acoustic wavefront propagation are likely to improve the resolution beyond what is observed in Fig. 4(b).

Fig. 4

Images of a chicken-breast specimen. (a) NCPAT image obtained with the following embedded objects (respective diameters in parenthesis): i, unpainted grayish metal wire (0.7 mm); ii, blood vessel phantom (0.5 mm); iii, blood vessel phantom (0.3 mm); iv, blood vessel phantom (0.7 mm); v, white painted metal wire (0.8 mm). (b) Corresponding NCUS image. All scales are in mm except for amplitude profiles (in arbitrary units).

3.2.

Images of a Calf Brain Specimen

Results obtained with a calf brain specimen are shown in Fig. 5. Similar embedded objects were used to test the imaging capability. Again, each embedded object is zoomed below the main NCPAT image [Fig. 5(a)] and NCUS image [Fig. 5(b)]. The corresponding horizontal profiles are also shown. As expected, the 0.8-mm-diameter white-painted metal wire (i) is only seen in the NCUS image while blood vessel phantoms with diameters of 0.7 mm (ii), 0.5 mm (iii), and 0.3 mm (iv) are only seen in the NCPAT image. In the NCUS image, the signal strength is good, but the lateral resolution is poor and is expected to be improved with better synthetic focusing obtained by using a three-dimensional (3-D) reconstruction algorithm, a two-dimensional (2-D) surface topography measurement, and the numerical modeling of the propagation of the generated ultrasonic wave. In this case, additional darker areas are also seen and may be first attributed to endogenous absorptions sites. However, due to the faster dehydration of the calf brain surface compared with the chicken breast surface, the reproducibility of results on calf brain was more difficult to verify. A validation method, such as x-ray computed tomography, would have been preferable, but this involves injecting a contrast agent in the blood prior to slaughtering the animal.

Fig. 5

Images of a calf brain specimen. (a) NCPAT image obtained with the following embedded objects (respective diameters in parenthesis): i, white-painted metal wire (0.8 mm); ii, blood vessel phantom (0.7 mm); iii, blood vessel phantom (0.5 mm); iv, blood vessel phantom (0.3 mm). (b) Corresponding NCUS image. All scales are in mm except for amplitude profiles (in arbitrary units).