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OPTOACOUSTIC DETECTION AND IMAGING OF THERMALLY OPTOACOUSTIC DETECTION AND IMAGING OF THERMALLY

DAMAGED EX VIVO TISSUESDAMAGED EX VIVO TISSUES

OPTOACOUSTIC DETECTION AND IMAGING OF THERMALLY OPTOACOUSTIC DETECTION AND IMAGING OF THERMALLY

DAMAGED EX VIVO TISSUESDAMAGED EX VIVO TISSUESTyler J Palmer1 and William M Whelan 1,2

1Physics, University of Prince Edward Island, 2 Biomedical Sciences, Atlantic Veterinary College, Charlottetown PEITyler J Palmer1 and William M Whelan 1,2

1Physics, University of Prince Edward Island, 2 Biomedical Sciences, Atlantic Veterinary College, Charlottetown PEI

The system consists of a Q-switched Nd:YAG pumped Ti:Sapphire laser delivering 1064 nm and775 nm light, respectively, through a bifurcated optical fiber bundle. Approximately 16 mJ of energywas delivered to the tissue over a 6 ns pulse duration with a 10 Hz repetition rate. An 8 elementannular array transducer with a 4 MHz central frequency was used for acoustic detection. Togetherthe elements have a focal point that is located 25 mm from the transducer surface. The signal fromeach of the 8 elements are processed to form a single A-line at each detection point.

To image each phantom, the box was mounted in the water bath with the mylar window 20 mmfrom the transducer surface. The transducer/laser fiber bundle scanned a vertical line in the centerof the window taking an average OA measurement (n=64) every 0.2 mm covering 17.2 mm (86 RF-lines). Each phantom was imaged at 775 nm and 1064 nm.

All data was exported to Matlab for analysis and individual RF-line inspection. A region of interest

Optoacoustic imaging is being investigated as a real-time technique for monitoring thermaltherapy tissue damage. This noninvasive imaging method uses pulsed laser energy to induce amechanical response in biological media which propagates through the surrounding medium asan acoustic wave and can be detected with a transducer.

In this study, samples of chicken breast tissue were held in a water bath at 50°C, 60°C or 70°Cfor 120 seconds. A reverse mode optoacoustic imaging system was used to acquire signals fromthe samples placed at a 5 mm depth within native chicken breast tissue. The illumination systemconsisted of a Nd:YAG pumped Ti:Sapphire laser delivering 1064 nm and 775 nm light,respectively, with a pulse duration of 6 ns. Acoustic signals were collected with an 8 elementannular array transducer operating at a 4 MHz central frequency.

The optoacoustic signals were analyzed using the Hilbert Transform and 2D images were

Figure 3. a) Frequency spectrum of representative RF-lines of damaged and native tissues from the70°°°°C 1064nm illumination scan. b) Normalized optoacoustic images for 1064nm illumination with lowpass (<2MHz) filter. Black boxes used for data in figure 4.

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All data was exported to Matlab for analysis and individual RF-line inspection. A region of interestwas chosen between 24-29 mm from the transducer (assuming a speed of sound of 1.54 mm/µs).A Hilbert transform was applied to each RF-line and images of the region of interest wereconstructed by mapping the normalized signal values to a color map. A Fast Fourier transform(FFT) was applied to the Hilbert transform over a region of interest. A comparison of the FFT plotsof the RF-lines “on” the heated sample and “off” helped to determine an appropriate data filter forimaging. The mean values and standard deviations of a 1x5 mm region within the expectedlocation of each thermally damaged sample were compared to get a qualitative analysis of thedetection and imaging abilities of the system.

The optoacoustic signals were analyzed using the Hilbert Transform and 2D images wereconstructed. Thermal damage was clearly visible in the images of the tissue samples heated to60°C and 70°C. When comparing the 50°C and 70°C samples, the average signal strengthincreased by factors of 1.6 and 2.5 for 775 nm and 1064 nm illuminations, respectively.Interestingly, differences in the signal frequency spectrum between native and damaged tissueswere observed below 2 MHz. Hence, the use of a low pass filter (<2 MHz) improved imagequality. This work demonstrates the potential of optoacoustic imaging to monitor tissue damagein thermal therapy applications.

Laser optoacoustic (OA) systems use laser pulses to generate pressure waves in a samplewhich are then detected with an ultrasonic transducer. Detection of the pressure waves over timeand in a 2D plane allows for 3D imaging of the irradiated sample. The generation of the pressurewaves depends on the light fluence, optical, thermal and mechanical properties of the sample,hence heterogeneities in the sample can be detected and imaged. The high resolution ofultrasonic signals combined with the high contrast due to light absorption has allowedoptoacoustic (alternatively, photoacoustic or thermoacoustic) systems to image breast cancer, ratbrain structure and function, tissue vasculature and tumor angiogenesis1. Other OA systemshave been used to detect changes in sample temperature2, to monitor blood oxygenation1 andfor tissue spectroscopy3. Laser OA has also been investigated as a potential monitoringtechnique for therapeutic heating4.

Thermal therapy is the application of heat to destroy tissue cells. Specifically, heating tissues totemperatures above ~ 55°C results in protein denaturation and cell death (coagulation)5. Thistreatment method has been under investigation for use as a minimally invasive technique toselectively eradicate benign tumors and cancer cells. One challenge associated with this methodis to ensure complete coagulation of the target tissue while avoiding damage to surroundinghealthy tissues. Laser OA has been proposed as a real-time noninvasive method for monitoring

Figure 1. Diagram of backward mode optoacoustic system showing the transducer/laser fiber bundleand tissue phantom in fiberglass box. a) native tissue b) thermally damaged tissue c) Mylar window.

pass (<2MHz) filter. Black boxes used for data in figure 4.

A Fast Fourier Transform was applied to the 70°C Hilbert Transform data. A significant increasein the frequency content of damaged tissue compared to native tissue below ~ 2 MHz wasobserved, as shown in figure 3a. Hence a low pass (<2 MHz) filter was applied and theresulting filtered images are shown in figure 3b. The filtered images appear less grainy andhave a concentrated region of higher signal values, clearly showing the location of the damagedtissues. This suggests that FFT analysis of optoacoustic signals may serve as a basis forchoosing a filter to help improve image quality.

Figure 4. Normalized optoacoustic signals (±±±± 1 standard deviation) from the regions of interest shownin figure 3b (n=650) as well as 775nm data from same location.

Figure 4 shows an increase in the optoacoustic signal strength by a factor of 1.6 with tissueheating from 50°C to 70°C, for 775 nm illuminations. For 1064 nm illuminations, the signalincreases by a factor of 2.5 over the same temperature range. This demonstrates that there is adifference in image contrast achieved for the two illumination wavelengths. It should be notedthat the maximum value in the 1064 nm 70°C scan was larger than the maximum value in the775 nm 70°C scan by a factor of 1.4 despite similar incident beam energies (difference of 0.4mJ). This suggests that the 1064 nm wavelength may be more suitable for detection ofthermally damaged tissues.

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This study demonstrates that thermally damaged tissues at a 5mm depth within native tissuescan be imaged by laser optoacoustic technology. The results indicate that optoacoustic imagingcan be sensitive to 10°C changes in a heating protocol and that there is an increase in signal byas much as a factor of 2.5 with tissue heating from 50°C to 70°C. This work also indicates thatlow pass filtering (< 2 MHz) can improve image quality.

healthy tissues. Laser OA has been proposed as a real-time noninvasive method for monitoringthermal tissue damage. The changes in the optical properties of tissues during coagulation havebeen documented6,7 and detected with an OA system by Larin et al.

In this study we investigate the ability of a laser optoacoustic system to detect and image thermaldamage in phantoms containing native and thermally damaged chicken breast tissue. We aim todetermine the sensitivity of the system to temperature-induced damage and to discuss a signalanalysis method for future use in the optoacoustic field.

1. Xu, Minghua and Wang, Lihong. Photoacoustic imaging in biomedicine. Rev. Sci. Instrum. 77 (2006).2. Pramanik, Manojit and Wang, Lihong. Thermoacoustic and photoacoustic sensing of temperature. J. Biomedical Optics, 14(5) (2009). 3. Kostli, Komel, et. al. Optoacoustic infrared spectroscopy of soft tissue. J. Applied Physics, 88(3) (2000). 4. Larin, Kirill, et. al. Monitoring of tissue coagulation during thermotherapy using optoacoustic technique. J. Phys. D: Appl. Phys. 38, 2645-2653 (2005).5. Diederich, Chris. Thermal ablation and high-temperature thermal therapy: Overview of technology and clinical implementation. Int. J. Hyperthermia. 21(8), 745-753 (2005). 6. Splinter R, et. al. Optical properties of normal, diseased, and laser photocoagulated myocardium at the Nd-Yag wavelength. Lasers Surg. Med. 11, 117-124 (1991).7. Essenpreis M, Thermally induced changes in optical properties of biological tissues. PhD Thesis University College London (1992)

Figure 2. a) Representative raw data RF-lines from damaged and native tissues (1064nm). b)Normalized optoacoustic images of the 1064nm illumination scans.

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The RF-lines in figure 2a are representative of measurement locations of both damaged (red) andnative (blue) tissues from the 1064 nm illumination scans. There is an increase and subsequentdecrease in the damaged tissue RF-lines from tissues damaged at 60°C and 70°C, however thereis no measurable difference between the two RF-lines in the 50°C scan. The maximum values andpeak to peak values increase with increasing heating temperature. The increase in signal strengthwith heating temperature is also evident in the optoacoustic images in figure 2b.

The location of the initial signal increase in the 60°C and 70°C images corresponds with thephysical location of the surface of the damaged samples. It is likely that the surface of thedamaged tissue is also responsible for the subsequent increase in signal strength in the 60°C and70°C scans. This is consistent with other published studies of optoacoustic detection of thermallydamaged tissues4. The apparent vertical size of the heated samples is larger than the actual sizeof 5mm. This is likely due to the transducer resolution of 2mm at the focal point.

Three tissue samples, 5 mm x 23 mm x 3 mm, were cut from a single chicken breast. They werethen individually covered with plastic wrap to prevent water contact. Samples were heated in aconstant temperature water bath of 50°C, 60°C or 70°C (±1°C) for 120 seconds. Uniform heatingwas assumed based on the sample size and the length of the heating time.

Two slabs of chicken breast tissue were cut to be ~30 mm x ~25 mm x 5 mm. These slabs wereused to sandwich each heated sample to simulate thermal damage at a 5 mm depth withinnative tissue.

A custom designed water tight fiberglass box was used to house the tissue phantoms (see figure1). The box had a removable front plate with a 20x20 mm, 25.4 µm thick optically transparentmylar window centered on the transducer axis. The box also contained a fiberglass plate whichwas adjusted to be 13 mm from the front window to hold the phantom in place. The optoacousticsystem consisted of a personal computer, laser, transducer and water bath. In the water baththere was a sample mount and a transducer/laser fiber bundle which could move by preciseincrements in the x-z plane (see figure 1).

a) b)