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1 March 2010 Saturation effect in functional photoacoustic imaging
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We investigate the saturation effect, which describes the violation of the linearity between the measured photoacoustic amplitude and the object's optical absorption coefficient in functional photoacoustic imaging when the optical absorption in the object increases. We model the optical energy deposition and photoacoustic signal generation and detection in a semi-infinite optical absorbing object. Experiments are carried out by measuring photoacoustic signals generated from an ink-filled plastic tube. The saturation effect is studied by varying the optical absorption coefficient in the model and the ink concentration in the photoacoustic experiments. By changing the center frequency of the ultrasonic detector, the requirement to minimize the saturation effect in functional photoacoustic imaging is established.



Photoacoustic (PA) imaging is a novel hybrid imaging technology that is able to image in great depth the optical energy deposition with high spatial resolution in biological tissue.1, 2, 3 Since optical energy deposition reflects the physiologically specific optical absorption contrast, PA imaging has been widely used to study various functional parameters of living biological tissue, especially hemoglobin oxygen saturation4, 5, 6, 7 (sO2) .

To measure sO2 , PA imaging depends on the difference in the molar extinction coefficient spectra of the two forms of hemoglobin, oxygenated hemoglobin (HbO2) and deoxygenated hemoglobin (HbR). When a blood vessel is irradiated by laser pulses with identical pulse energy at different optical wavelengths, PA signals are generated with different amplitudes due to the optical-wavelength-dependent molar extinction coefficients. Therefore, PA imaging performs spectral measurement and treats HbR and HbO2 as the dominant optical absorbers at each wavelength (λi) to calculate their relative concentrations ([HbR] and [HbO2] , respectively) based on a least squares fit.5 The sO2 is further calculated by [HbO2]([HbO2]+[HbR]) .

The prerequisite for the aforementioned spectral measurement and inverse calculation (least-squares fit) is that the detected PA signals from blood are linearly related with the optical absorption coefficients of hemoglobin at λi , which denotes as the absorption-dependent PA detection. However, when the center frequency of the ultrasonic detector is low with respect to the corresponding optical absorption coefficient,8 the detected PA signal will not linearly reflect the optical absorption change in blood. In an extreme case, the detected PA signal may rely on only the incident optical energy regardless of the optical absorption coefficient of blood. This phenomenon is referred to as the saturation effect.9 When saturation effect occurs, the inverse calculation of the [Hb] and [HbO2] becomes inaccurate.

In the in vivo PA imaging of sO2 , the required maximum imaging depth for different applications varies from a few hundred micrometers to several millimeters. In general, increased maximum imaging depth requires the ultrasonic detector to have a lower center frequency because of the wavelength-dependent ultrasonic attenuation in tissue.10 Due to the saturation effect, however, a threshold of the ultrasonic center frequency could exist. Using an ultrasonic detector with a center frequency below such a threshold will lead to saturation in the PA signal detection and present an incorrect inverse calculation of sO2 with an error greater than a preset value.

Here, we report both the theoretical and experimental studies of the saturation effect in functional PA imaging. In our experiments, the change of the optical absorption coefficient was achieved by varying the concentration of the optical absorbing material rather than the optical wavelength. The saturation effect discussed here is limited to the nonlinear dependency of the measured PA amplitude with the optical energy deposition under the adiabatic heating condition induced by short laser pulses. Other nonlinear effects caused by, for example, heat diffusion11 are excluded.


Methods and Materials

We used a semi-infinite homogeneous optical absorbing object, as shown in Fig. 1 , to model the optical energy deposition as well as the PA wave generation and detection. This model is valid as an approximation for blood vessels when the optical penetration depth is less than one-tenth of the vessel diameter. Within the spectral range of 520to600nm , the optical absorption coefficient of the whole blood (considering the sO2 to be 95%) is12 20to296cm1 , corresponding to a penetration depth of 34to491μm . In our model, the laser illumination is along the z axis and the ultrasonic detector is placed directly above the illuminated area.

Fig. 1

Semi-infinite optical absorbing object to model optical energy deposition and PA signal generation.


The depth-resolved optical fluence distribution F(z) inside the object is described as

Eq. 1

where F0 (in joules per squared centimeter) is the optical fluence on the surface of the object, and μa (in inverse centimeters) is the optical absorption coefficient of the medium. Hence, the induced depth-resolved PA waves can be described as

Eq. 2

where k1 is the conversion coefficient of the PA effect (dimensionless), which is the result of the Grueneisen parameter (dimensionless) multiplied by the percentage of optical energy deposition that is converted to heat.13

As the peak amplitude of the detected PA signals is usually taken as a measure of the optical energy deposition in a vessel, here we estimate the peak amplitude of the detected PA signal for Eq. 2. If an ultrasonic detector has an infinitely high axial resolution, the detected peak PA amplitude will be αF0k1μa , where α is the coefficient of the ultrasonic detection system that converts pressure to voltage. For an ultrasonic detection system that has only finite axial resolution, the detected PA signal is an integration of the PA waves within a specific axial pixel. For a given wideband ultrasonic detector at a cut-off wavelength of λc , the axial resolution in PA imaging can be approximated14 as k2λc , where k2 is a dimensionless coefficient between 0.8 to 0.92. Thus, the detected PA peak amplitude can be approximated as

Eq. 3

APA(μa)=α0k2λcF0k1μaexp(μaz)dz =αF0k1[1exp(μak2λc)].

The experiments were carried out by measuring PA peak-to-peak amplitudes detected from an ink (Fiesta Red, Private Reserve Ink, The Ink Flow, San Antonio, Texas ) filled plastic tube (Tygon® S-54-HL Microbore Tubing 56464, United States Plastic Corp., Lima, Ohio ). The measured optical absorption spectrum of the ink is shown in Fig. 2 . The optical wavelength used for the PA measurement was 523nm . The corresponding optical absorption coefficient was 1028cm1 . The tube has an inner diameter of 0.5mm and an outer diameter of 1.5mm .

Fig. 2

Measured optical absorption spectrum of the red ink at 100% concentration.


A schematic diagram of the experimental setup is shown in Fig. 3 . Laser light at 523nm was generated by a Nd:YLF laser (INNOSLAB; IS8II-E; EdgeWave, Germany ). The pulse width was 6ns , the pulse repetition was set to 100Hz , and the laser beam diameter was 10mm . The light coming out of the laser first passed through a beam sampler, then a right-angle prism (PS908; Thorlabs, Newton, New Jersey ), and finally a neutral density filter (NE10A; optical density: 1.0; Thorlabs, Newton, New Jersey ) before it illuminated the tube. The light was attenuated because of better pulse energy stability when the laser system is working at a higher driving current.

Fig. 3

Schematic of the experimental setup.


The PA waves were detected by a high-frequency, unfocused ultrasonic transducer and amplified by a wideband, lownoise preamplifier (ZFL500LN+, Mini-Circuits, Brooklyn, New York ) for 28dB . The amplified PA signals were then digitized and stored by an oscilloscope (TDS5034B, Tektronix, Beaverton, Oregon ). A fast photodiode (DET10A, Thorlabs, Newton, New Jersey ) detected the sampled laser pulses to generate triggers for data acquisition. Each PA signal was averaged 600 times to minimize the influence from the pulse energy instability.

Three ultrasonic detectors were used in the experiments. Their respective center frequencies and bandwidths (6dB) were 10 and 9MHz (V312, Olympus, Waltham, Massachusetts ), 20 and 10MHz (V317, Olympus, Waltham, Massachusetts ), and 40 and 15MHz [custom-built by the National Institutes of Health (NIH) Resource Center for Medical Ultrasonic Transducer Technology,15, 16 The University of Southern California, Los Angeles]. Their corresponding cut-off wavelengths were 105, 60, and 32μm . Each transducer was positioned independently to reach its maximum PA amplitude.


Results and Discussions

Figure 4 gives the theoretical calculations of PA(z,μa) and APA(μa) . Figure 4 shows the normalized depth-resolved PA waves along the z axis in the semi-infinite object calculated from Eq. 2. The μa was varied from 70to630cm1 with a step size of 140cm1 . For simplicity, all the PA(z,μa) were normalized by F0k1μamax , where μamax is the maximum value of μa used in simulation study, which was 720cm1 . As we can see in Fig. 4, higher μa induces larger PA pressure peak amplitude under the same optical illumination, but the depth-resolved signal decays faster than those with lower μa values.

Fig. 4

Theoretical study of the PA pressure and the detected PA amplitude: (a) depth-resolved PA pressure induced by a delta laser pulse illumination in media with different optical absorption coefficients and (b) correlation of the detected PA peak amplitude with μa detected by different ultrasonic detectors.


When an ultrasonic detector measures the induced PA pressures for anatomical or functional imaging purposes, a time-resolved PA signal will be acquired. In most situations, the peak-to-peak amplitude of the detected PA signal is used to measure functional parameters such as sO2 . Ideally, the detected PA signal is anticipated to be identical with the induced PA pressures inside the sample; however, because ultrasonic detection is band-limited and thus has only limited axial resolution, the detected PA signal must be integrated within an ultrasonically resolvable depth. Such resolvable depth can be considered as the axial resolution and is related with the ultrasonic center frequency and bandwidth.10, 14 In our work, the axial resolution is approximated by k2λc .

Figure 4 shows the calculated APA(μa) for the three ultrasonic detectors used in the experiments. In the simulation, k2 was set to 0.92 and the APA(μa) for each ultrasonic center frequency was normalized independently. As we can observe in Fig. 4, when μa is low, APA(μa) increases proportional to μa for all three center frequencies; however, when μa is higher than a certain value, the correlation between μa and APA(μa) loses its linearity and eventually reaches a plateau, which demonstrates the saturation effect. Moreover, the nonlinearity becomes more severe when the ultrasonic center frequency lowers. In Fig. 4, the PA signal detected by a 10-MHz ultrasonic detector was saturated at a μa value lower than those detected by 20- and 40-MHz ultrasonic detectors. If we consider the μa value corresponding with the 1e decay of APA(μa) as the “turn point” of the saturation, the linear range for a given λc can be found from the simulation results. As seen in Fig. 4, the horizontal line corresponding with the value of 1e1 , where the left side of the cross point with each APA(μa) representing the preferred linear measuring range. It is important to understand that such a transition at the 1e “turn point” between linearity and saturation is not abrupt.

The saturation effect limits the selection of the ultrasonic detector’s center frequency and bandwidth in functional PA imaging since the inverse calculation of sO2 requires the APA(μa) to be linearly associated with optical absorption. Due to the dramatic variation in the molar excitation coefficients of Hb and HbO2 within the spectral range of 520to600nm , the PA measurement may be saturated within a certain spectral range while remaining linear within others. The saturation effect was further observed by measuring PA signals generated from the ink phantoms. The same normalization scheme as that used for Fig. 4 was used in the experiments.

Figure 5 shows the change of the detected PA amplitude using the 20-MHz ultrasonic detector when the ink concentration increases. The temporal profiles of several detected PA waves are also shown in the Fig. 5. Similar temporal shapes can be observed for PA waves within both the linear and saturated ranges. As shown in Fig. 5, the experimental data acquired from three ultrasonic detectors demonstrate the same trend as that in the simulation results, where the saturation starts much earlier for the 10-MHz transducer than transducers with higher center frequencies.

Fig. 5

Experimental study of the saturation effect in ink phantoms: (a) detected PA amplitudes acquired from various ink concentrations using the 20-MHz detector, the temporal profiles of the PA signals measured at ink concentrations of 5, 15, 28, 36, and 50% are plotted at their corresponding μa values, and (b) correlation between the detected PA peak amplitude μa of the ink phantom for different ultrasonic center frequencies. P-P, peak-to-peak.


We further fit the experimental data to verify k2 . The target fitting function was simplified from Eq. 3 to be a-b (ecμa) , where a , b , and c are unknown constants. After fitting was conducted for all three ultrasonic transducers, we acquired a=0.98±0.016 , b=0.82±0.090 , and c=0.0076±0.0018 . Then k2 was subsequently calculated to be 1.15±0.10 . Compared with the values reported in Ref. 14, our value is only slightly higher, possibly due to experimental noises and different optical-ultrasonic geometry. Both the simulation and experimental results indicate that the ultrasonic detector used for functional PA imaging should satisfy k2λcμa1 to avoid saturation. In this situation, Eq. 3 becomes APA(μa)=αF0k1k2λcμa by ignoring high-order terms in the Taylor expansion.



We have used a semi-infinite homogeneous optical absorbing medium to model laser-induced, depth-resolved PA pressure and the resolution-limited detected PA amplitude. The saturation effect was simulated by varying the optical absorption coefficient of the medium. The saturation effect was further observed in ink phantoms using three ultrasonic detectors with center frequencies of 10, 20, and 40MHz . The experimental results agreed with the theoretical studies and the fitted constant agrees with the reported theoretical estimation.


This work was supported in part by start-up funding from the University of Wisconsin-Milwaukee, a grant from The Lynde and Harry Bradley Foundation, and financial support from the Chinese Ministry of Education to J. Wang.



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©(2010) Society of Photo-Optical Instrumentation Engineers (SPIE)
Jing Wang, Tan Liu, Shuliang Jiao, Ruimin Chen, Qifa Zhou, K. Kirk Shung, Lihong V. Wang, and Hao F. Zhang "Saturation effect in functional photoacoustic imaging," Journal of Biomedical Optics 15(2), 021317 (1 March 2010).
Published: 1 March 2010

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