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In this paper a new generation of optical-resolution photoacoustic microscopy (OR-PAM) with a wide range of potential clinical applications is demonstrated. Using fast scanning mirrors, an image guide with 30,000 fiber pixels, a refocusing lens and a unique probe we managed to reduce the footprint of an OR-PAM system from a stationary table-top system to a portable, 4cm by 6cm, probe weighing ~500g tethered to a scanning unit. The phantom studies show that the handheld optical-resolution photoacoustic microscope is able to image with ~7μm resolution. For in vivo studies images of the microvasculature in a Swiss Webster mouse ear are shown. The compact, flexible nature of the proposed design and the small footprint of the apparatus increase the usability of OR-PAM for potential clinical applications such as in dermatology.
©2011 Optical Society of America
1. Introduction
Optical-resolution photoacoustic microscopy (OR-PAM) is an imaging technology providing high micron-scale lateral spatial resolution to visualize superficial structures in vivo with optical-absorption contrast [1]. The term “optical resolution” is used since the lateral resolution of the system is defined by the optically-focused spot size, which is limited by the diffraction-limit. The axial resolution of the system is still inversely related to the bandwidth of transducer, typical of other photoacoustic microscopy (PAM) systems.
Due to limitations of light transport, OR-PAM is used for imaging superficial structures to depths of about 1 mm in tissue [2,3]. OR-PAM is able to image capillary networks and quantify morphological and functional parameters such as number of vessels, diameter and length of the vessels, total hemoglobin concentration and hemoglobin oxygen saturation [2,4,5].
Since cancer cells need to consume more oxygen and nutrients to grow compared to other tissues [5,6], they signal the onset of angiogenesis and hence have plenty of contrast-providing blood vessels, making OR-PAM a useful tool for tumor imaging. Real-time OR-PAM may have a wide range of potential applications for clinical practice [7,8]. Recently a real-time B-scan OR-PAM system was demonstrated [9] with scanning speed up to 40 Hz over 1-mm range or 20 Hz over 9-mm range. In order to develop a real-time OR-PAM system a high laser pulse repetition rate (PRR) and fast data acquisition and display system are required. Laser systems such as flash lamp-pumped laser systems (PRR ~up to100 Hz) and diode- pumped solid-state Q-switched lasers (PRR ~up to few kHz) typically do not have high enough repetition rates for real-time C-scan imaging [7].
Currently laser systems with repetition rates greater than 100 kHz, adequate pulse durations, and energies for real-time OR-PAM have not yet been widely explored for photoacoustic imaging [7]. We introduce a high repetition rate, inexpensive, compact laser source for realizing high frame rate photoacoustic imaging [7,8].
Another limitation of current OR-PAM systems is their lack of flexibility. The present OR-PAM systems are mostly mounted as a table-top device with a large footprint. However, recently we introduced label free optical-resolution photoacoustic micro-endoscopy to enable access to internal body cavities [10]. Thus far, however, an external ultrasound transducer is required. To extend the potential range of applications, for the first time, we demonstrate a handheld real-time optical-resolution photoacoustic microscope (HH-OR-PAM) with a 4 cm by 6 cm footprint and weight of less than 500 g using image guide fibers and a unique fiber laser system. Additionally, the capability of HH-OR-PAM is demonstrated by phantom and in vivo (microvasculature in a Swiss Webster mouse ear) studies based on a customized nanosecond-pulsed tunable fiber laser with repetition rates of up to 600 kHz, enabling near real-time C-scan and volumetric imaging.
2. System description
Figure 1a shows the experimental setup of the laser scanning HH-OR-PAM. A pulsed fiber laser (YLP-G, IPG Photonics Corporation.) with second-harmonic wavelength of 532nm is able to provide tunable average output power up to 13W and pulse energy up to 20µJ. A glass slide is used to deflect a small amount of light into a photodiode. The photodiode is used to trigger the high speed data acquisition card (Gage card CS8289). The laser beam passes through a 2D galvanometer scanning mirror system (6230H, Cambridge Technology Inc.). The mirrors are controlled by separate drivers. These drivers can run with an analog control signal or a digital waveform depending on the study requirement.
Fig. 1 (a) Experimental setup of our HH-OR-PAM imaging system employing a high repetition rate diode- pumped pulsed Ytterbium fiber laser with up to 600 kHz pulse repetition rate and pulse widths of 1ns. FLD: fiber laser driver, Yb: Ytterbium, M: mirror, Gs: glass, PD: Photodiode, FG1: function generator channel 1, FG2: function generator channel 2, DX: X axis mirror driver, DY: Y axis mirror driver, OL: objective lens, A: amplifier (Olympus 5900PR). (b) The structure of the handheld probe. The light at the end of the fiber is refocused using a pair of glass aspheric lenses with 350 to 700nm AR Coating. Then the light passes through an oblique 10-mm fused silica prism. The photoacoustic signals directed upward to the prism’s diagonal will be deflected to a focused transducer (f = 19mm). AL: glass aspheric lenses, UST: ultrasound transducer, P: prism, IMF: index-matching fluid. (c) The prototype of the handheld probe. AL: glass aspheric lenses, UST: ultrasound transducer, P: prism, IMF: index-matching fluid.
In this experiment the mirrors were driven with sine waves generated by a two channel function generator (Tektronix AFG3022B). The frequency of the control signal determines the speed of scanning, while the peak-to-peak amplitude determines the angle of the scanning mirrors and eventually the field of view. In this experiment, the fast and slow scanning rates are fixed at 400 Hz and 1 Hz respectively, providing 2 C-scan frames per second. Using an objective lens (f = 18 mm), the scanning light via mirrors was coupled to the 1m-long image guide fiber with 30,000 individual single mode fiber elements (Sumitomo IGN-08/30). The field of view is limited by the diameter of the image guide (800µm). The end of the fiber is connected to the handheld probe.
Figure 1b shows the structure of the handheld probe. The light at the end of the fiber is refocused using a pair of aspheric glass lenses with 350 to 700nm AR Coating (Thorlabs 350260 and 352220). Then the light passes through an oblique 10-mm fused silica prism. The photoacoustic signals directed upward to the prism’s diagonal will be deflected to a 10-MHz focused ultrasound transducer (19-mm focus, 6-mm active element, f# = 3.17, CD International Inc). Calculations show that most of the acoustic energy is preserved at the prism interface within 45 ± 13.4 degrees within the angular acceptance of the transducer [7]. Optical index-matching fluid is used in order to allow the top-down laser illumination to be guided to the imaging target without optical refractive path variation. Acoustic attenuation loss in index-matching fluid is only slightly greater than water [11]; therefore, it is also used for ultrasonic coupling.
3. Result and discussion
In our experiments the fiber laser is utilized to produce 100mW; however, considering coupling efficiency of the image guide and other optical losses, at the end of the handheld probe the power is reduced to 45mW. The PRR is set to 160 kHz, therefore the pulse energy on the sample is calculated as ~0.3 µJ. This power is fixed for the entire experiment. The high speed and slow speed scanning mirror are driven by 400Hz and 1Hz analog sine wave respectively with 400mV peak-to-peak voltage.
In order to show the ability of the proposed design, different phantom targets have been imaged as shown in Fig. 2 . Figure 2a and 2b shows a network of carbon fibers with ~7.5µm diameter which are as small as capillary sized blood vessels. Figure 2c shows an image of a human hair with a diameter of ~100µm. For the in vivo studies, we employed the HH-OR-PAM to image the microvasculature in the ear of a Swiss Webster mouse. All the experimental procedures were carried out in conformity with the laboratory animal protocol approved by the University of Alberta Animal Use and Care Committee. Authors are also trained and certified in order to use mice in the research work. During the imaging session the animal was anaesthetized using a breathing anesthesia system (E-Z Anesthesia, Euthanex Corp.). Figure 3 shows images of the microvasculature in the ear of a Swiss Webster mouse. A 2D Hessian-based Frangi Vesselness filter [12] was used to filter C-scan maximum amplitude projection (MAP) images to preferentially select tubular structures while rejecting noise.
Fig. 2 (a) and (b) network of carbon fibers with ~7.5µm diameter which are as small as capillary sized blood vessels. (c) Human hair with a diameter of ~100µm.
Measuring the average signal amplitude and the standard deviation of the noise, the signal-to-noise ratio for the carbon fiber is calculated as ~22dB. Fitting individual carbon fiber signal amplitude to a Gaussian function, the full-width-half-maximum (FWHM) is calculated approximately 8.7µm (Fig. 4 ). The measured FWHM shown in Fig. 4 is partly due to the 7.5-μm width of the carbon fiber itself [7]. In order to measure the true optical lateral resolution, the convolution of a 2-D Gaussian beam with a carbon fiber is simulated for a range of spot-sizes, and compared with measurements [7]. Using this technique, the optical lateral resolution is estimated as 7µm. As mentioned the fast and slow speed mirror scanning rates are set at 400 Hz and 1 Hz respectively, with the laser repetition-rate set at 160 kHz. This corresponds to 2 volumetric-scans (composed of 400 B-scans per volume) per second. C-scans are formed from maximum-intensity projections of A-scan lines and contain 200 non-uniformly-spaced pixels in the X-direction and and 400 in the Y-direction, respectively. Higher C-scan frame rates are easily attainable; however, the field-of-view must be sacrificed if the resolution is to be maintained.
For a 400 by 400µm field of view (FOV) the average step size for the X and Y directions are calculated as 2 and 1µm, respectively. Since the scanning trajectory of the laser scanning system employs sine wave driving waveforms, the step size at the middle of scanning trajectory can be larger compare to its edge. Using Eq. (1) and Eq. (2) the worst case step sizes for the X and Y directions are calculated as ~3µm and 2µm respectively.
where FR is Fast scanning rate, SR is slow scanning rate, FOVx is field of view in X direction, FOVy is field of view in Y direction and PPR is pulse repetition rate. The lateral spatial resolution (FWHM) is calculated as ~7μm, or Gaussian sigma parameter, σ = FWHM / (2√2ln2) ~3μm. Considering the total area of the fiber bundle with 30,000 individual single mode fibers, we obtained the diameter of each strand to be about 4μm which is close to the measured lateral resolution. This shows that the system is almost at optimal resolution. For the in vivo study, assuming that the depth of the laser focus is ~200μm below the tissue surface and that the numerical aperture of the lens is about 0.16, the calculated laser surface fluence is ~8mJ/cm2 which is lower than the American National Standard Institute (ANSI) safety limit (20 mJ/cm2) [13].Fiber-based imaging systems have been demonstrated before [14–16]. Such systems also rely on image-guide fibers, miniature fiber bundles capable of transmitting images. Many such systems require the use of an exogenous fluorescent dye or contrast agent. HH-OR-PAM would provide the ability to image microvessels, and with improved resolution, potentially cells and subcellular structures similar to fluorescence (confocal) micro-endoscopes, but without necessarily requiring an exogenous contrast agent.
We anticipate that the HH-OR-PAM system will open up new possibilities for clinical and pre-clinical uses, such as functional brain mapping, cancer imaging and detection, assessing completeness of melanoma resection, imaging of angiogenesis to assess therapeutic efficacy, etc. The proposed label-free system will have potential for translational research because it will be compact and potentially portable, real-time, cost-effective. It will permit clinical imaging of different parts of body that were previously inaccessible.
4. Conclusion
The development of real-time HH-OR-PAM system with a ~4 by 6-cm footprint and weight of ~500g is demonstrated. The probe consists of an image-guide fiber bundle, a pair of glass aspheric lenses, a prism, index-matching fluid and a focused ultrasound transducer. The image guide is made of 30,000 individual fibers in an 800μm bundle. A fiber laser producing ~1 ns pulses with tunable energy up to 20μJ at a wavelength of 532 nm and a tunable repetition rate up to 600kHz was coupled into the 1m-long fiber using an objective lens focused on the fiber input and the beam was scanned across the fiber tip using a high speed mirror galvanometer. Both in vivo and in vitro images were demonstrated. The resolution study shows ~7µm lateral resolution. The proposed setup maintains all the powerful properties of previous OR-PAM system. Therefore, this system will improve the usability of OR-PAM and it could have clinical significance for a number of applications.
Acknowledgment
We gratefully acknowledge funding from NSERC (355544-2008, 375340-2009, STPGP 396444), Terry- Fox Foundation and the Canadian Cancer Society (TFF 019237, TFF 019240, CCS 2011-700718), the Alberta Cancer Research Institute (ACB 23728), the Canada Foundation for Innovation, Leaders Opportunity Fund (18472), Alberta Advanced Education & Technology, Small Equipment Grants Program (URSI09007SEG), Microsystems Technology Research Initiative (MSTRI RES0003166), University of Alberta Startup Funds, and Alberta Ingenuity/Alberta Innovates scholarships for graduate and undergraduate students.
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