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A novel prototype intraoperative system combining positron detection and optical coherence tomography (OCT) imaging has been developed for early ovarian cancer detection. The probe employs eight plastic scintillating fiber tips for preferential detection of local positron activity surrounding a central scanning OCT fiber providing volumetric imaging of tissue structure in regions of high radiotracer uptake. Characterization measurements of positron sensitivity, spatial response, and position mapping are presented for Tl204/Cs137 sources as well as 18F-FDG. In conjunction with co-registered frequency domain OCT measurements the results demonstrate the potential for a miniaturized laparoscopic probe offering simultaneous functional localization and structural imaging for improved early cancer detection.
©2009 Optical Society of America
1. Introduction
Positron emission tomography (PET), using 18F–fluoro-2-deoxy-d-glucose (18F-FDG) as a tracer, can detect malignant cancers with altered glucose metabolism and has been used for the assessment of lymph node involvement [1], evaluation of pretreatment staging and treatment response [2, 3] and detection of cancer recurrence. It also holds promise in more accurate evaluation of recurrent or residual [4, 5] than those morphologic modalities. However, it has limited value in lesion localization in early stages of ovarian cancer because of the difficulty in distinguishing between the signal from early-stage cancers and the background uptake signals coming from normal tissue [6]. Thus there is a substantial need for more efficacious imaging technologies. The ability to target pre-neoplastic or early neoplastic changes in the ovary with an imaging modality that required only a minimally invasive surgical procedure would greatly enhance care for women at risk for ovarian cancer.
Compared with external 18F-FDG PET imaging, intraoperative approaches have the significant advantage of detecting localized early-stage cancers with small amounts of radioactivity. The mean tissue path of gamma radiation is several centimeters for typical gamma energies of a few hundred keV. As a result, the gamma signal from a small tumor is contaminated by the signal from the surrounding normal tissue. Intraoperative probes that are selectively sensitive to short-range beta radiation have the significant advantage of high tumor-background contrast and therefore high sensitivity of detecting early-stage cancers. The tissue path length for beta particles is about 2.3 mm and it corresponds well with targeted internal organs, such as ovary, colon, and cardiac vessels before these particles interact with electrons to produce annihilating gamma photons. A number of investigators are developing intraoperative or intravascular radiation detection systems for early detection of cancers or atherosclerosis [7–16]. Many of the detectors proposed for use in beta-sensitive probes have utilized plastic scintillators [12, 13, 17]. Due to the low effective atomic number of this material, the detection efficiency for most background photons is small. Daghighian et al., for example, have developed a dual plastic scintillation probe in which background gamma events are detected and subtracted from the probe signal [8]. The catheter designed by Janecek et al. [12] employs 0.5 mm diameter scintillating fibers and has demonstrated acceptable sensitivity for detection of 18F of >400 cps/μCi at 1mm from the detector. While these groups are developing radiation detectors for intraoperative or intravascular use, none of these efforts have linked this type of functional intraoperative radiation detection with a device that allows simultaneous high-resolution structural imaging. Because early-stage cancers are usually associated with pre-neoplastic changes, an imaging modality that may provide composition and structure of the lesion offers a great advantage in detecting cancers at an earlier and more treatable stage. Optical Coherence Tomography has shown potential to evaluate early structural changes in the ovary.
Optical coherence tomography (OCT) is an emerging high resolution imaging technique [18]. OCT measures backscattered light, generated from an infrared light source directed to the tissues being examined. OCT typically obtains a resolution capability of 5 to 10 μm and a depth of penetration of 1–2 mm and has been used to image tissues in the body that can be accessed either directly or via endoscope or catheter, including the GI tract [19, 20], eye [21], ovary [22, 23], and coronary blood vessels [24, 25]. The morphological features of pre-neoplastic or early neoplastic changes have prompted development of this high-resolution imaging modality for early-stage ovarian cancer detection [26]. OCT is sensitive to changes in collagen that correlate with changes in collagen seen as malignancy develops [26]. OCT can also detect areas of necrosis and blood vessels that are indicative of an underlying abnormality in the tissue not detected by the surgeon.
The intraoperative scintillating probes provide limited spatial resolution and can, at most, identify activities about the size of fiber tips (approximately 1 mm). However, the scintillating probes are highly sensitive to targeted radiolabeled tracers. In our previous publication, we have introduced the first prototype OCT/scintillating device to demonstrate the principle of co-registration [27] and to perform ex vivo validation using rabbit atherosclerotic models, which have high 18F-FDG lesion-to-background contrast due to high metabolism [28]. The scintillator used in the reported studies was a single fiber that provided the detection curve. In this paper, we introduce a novel hybrid intraoperative imaging device, which integrates an OCT imaging probe and multiple scintillating fibers for mapping the activities of the positron activity. This hybrid imager has potential applications in ovarian cancer screening during minimally invasive endoscope procedures. The device is intended to interrogate a large surface of the ovary through a laparoscopic port to detect local 18F-FDG uptake using the scintillating probe and to image local morphological changes using the OCT probe.
During surgery, the positron detectors would locate regions of high radiotracer uptake due to high metabolic activity (as with 18F-FDG) or specific molecular markers (bioconjugated tracer) that might be correlated with early malignant changes. The OCT images would then be used for evaluation of tissue changes at the cellular or multi-cellular level such as collagen restructuring.
2. System design
2.1 Hybrid intraoperative probe
Figure 1 depicts the design of the hybrid scintillating/OCT system. Two rows of scintillating detectors provide coarse localization and detection of positron activity while a scanning OCT catheter provides high-resolution, co-registered structural imaging of the tissue through a central channel. Although the initial prototypes demonstrating the concept described below are larger, future versions of the probe will be miniaturized for insertion through a standard 5 mm laparoscopic port.
The nuclear detection section of the probe consists of eight positron detectors arranged in a 2 × 4 configuration straddling the OCT channel along with one gamma detector on each side for background estimation and subtraction (Fig. 1(b)). Each detector consists of a 1 mm diameter × 3 mm long plastic scintillator fiber tip (BCF-12, Saint-Gobain Crystal) epoxied to the same diameter, 1.2 m long optical fiber (BCF-98). The sides of each scintillator tip are coated with reflective paint to enhance light collection. In the gamma detectors, a 0.25 mm thick Cu plate shields the tips from positron events. All fibers are sheathed with flexible stainless steel braid-reinforced medical tubing (New England Catheter) to just beyond the tip/fiber junction for fiber protection and light shielding. The spacing of the scintillator tips along the longitudinal direction was 1.75 mm with a lateral (row-to-row) spacing of 4.25 mm. This configuration provides an imaging area of approximately 5 mm (lateral) × 6 mm (longitudinal).
The OCT imaging portion consists of a side-viewing catheter based upon a conventional SMF-28 single mode fiber with integrated graded-index imaging lens and 45-degree deflecting prism at the distal end. Two motor assemblies at the proximal end of the fiber provide simultaneous translation and rotation of the catheter to generate a sequence of 150-degree sector cross-sectional scans of the ovarian tissue at a defined longitudinal spacing (Fig. 1(a)). For the images presented the rotation speed was 4 revolutions per minute (RPM) although rates of up to 10 rotations per second are possible with the system. The linear pullback rate was stepper motor limited at 0.2 mm/sec. Due to the high sensitivity of the scintillation fibers necessary to detect the weak signals from nuclear events, OCT scanning is performed only after a high positron activity region has been identified.
The positron detection fibers are separated from the OCT fiber and input into a multiple channel detection system for scintillation counting. The OCT fiber connected to the sample arm connection of a high-resolution fiber-based frequency-domain OCT system.
Fig. 1. The prototype hybrid intraoperative probe system. (a) The single-mode OCT fiber is translated and rotated from the proximal end of the catheter and pulled out gradually during the scanning. (b) Tissue surface view of probe illustrating location and spacing of 1mm scintillating fiber tips alongside the central OCT slot. (c) Photograph of system.
2.2 FD-OCT system
The Fourier-Domain OCT system (Fig. 2(a)) features a 110 nm bandwidth swept-source (HSL-2000, Santec Corp., Japan) operating at 1310 nm providing 12 μm depth resolution (air) imaging with a 20 kHz scan rate suitable for in vivo use. At the working depth of 2 mm from the surface of the prism, the lateral resolution is 19 μm. For the reduced rotation rate used for the initial demonstrations, the scan rate was lowered to 250 Hz using a divisor circuit. The resulting images therefore contained 2110 A-lines corresponding to a lateral over-sampling of a factor of over 2 at the working distance (180-degree). The reflected sample signals and transmitted reference beams are combined via a 3dB coupler, converted to electrical signals with balanced detection (ThorLabs PDB120C), filtered with a 20 MHz anti-alias filter, and digitized with a 50 MHz analog-to-digital converter (CS8325, GaGe Applied).
During image processing, a wavelength-dependent amplitude correction to the raw data is applied to account for the optical power variation during A-line scans. The scaled measurement data is interpolated to a uniform frequency grid before inverse FFT transformation. The image is log-compressed to produce a 1760 A-line OCT image with 55 dB dynamic range. A 5×5 pixel median filter was applied to final images to improve contrast. Total processing and display time was 3 seconds per image; this duration can be reduced to < 1 second per cross-section for real-time feedback with software optimization and reducing the A-lines to 1024. An example of ex vivo OCT image of normal human ovary is shown in Fig. 2 (b). The patient was a 38-year-old high risk premenopausal woman.
Fig. 2. (a). Swept-source Fourier-Domain OCT system. (b) Ex vivo OCT image of human ovary epithelium. Comparing the OCT image with the histology, we see inclusion cysts (cysts in ovarian surface epithelium) about 200 μm deep (white arrows). Adjacent, we see the edge of a follicle (blue arrow) about 400–600 μm deep. Higher intensity regions indicate rich collagen content (red arrows) [26].
2.3 Positron detection system
Figure 3 depicts the block diagram of the detection system consisting of a photomultiplier, front-end electronics, and sampling-based multi-channel analyzer circuitry. The probe fibers are terminated on a Hamamatsu R8900U position sensitive PMT (PSPMT) for detection and imaging. The PSPMT and front-end electronics assembly is sealed in a light-tight box for experimentation to reduce stray photon background. The PMT provides 6 X and 6 Y readout anodes for position determination which, in turn, are amplified and connected to a resistor divider network to condense to four outputs: X1, X2 and Y1 and Y2. Standard Anger logic was used to calculate the event position and the total event energy [29].
The BCF-12 blue-emitting scintillation fiber in our system has decay time of 3.2 ns, which is faster than common scintillators such as NaI (230 ns), BGO (300 ns) and LSO (47 ns). The fast decay times necessitated modifications to the detection circuitry and approach to enable accurate energy and image position determination. First, the inter-anode resistances were reduced to 12.5 Ω [30] from values of 1000Ω to 1MΩ in traditional radiation counting. For the higher frequencies present in the short scintillation pulses, package, PMT, and cabling parasitics produce strong frequency dependencies in the splitting ratio between the complementary resistor network output paths. These parasitics result in a pulse-dependent splitting ratio and increased position uncertainty. The lower resistances yielded tighter localization in the Anger logic positioning diagrams at the expense of reduced Anger logic ratio separation between fiber locations.
The second design consideration was the adoption of a sampling-based approach for determination of peak pulse voltage or total pulse energy. Conventional peak-hold detection circuits are limited to pulse durations of approximately 200 ns or larger. Improvements in high-speed analog-to-digital converter technology, however, have made it possible to directly sample signals as short as 5–10 ns. To further improve the fidelity of pulse detection, custom shaping amplifier stages that filter the pulses to approximately 15 ns duration independently amplified the four resistor network outputs. A four-channel, 8-bit GS/s data acquisition system (Octopus CS22G8, Gage-Applied, Inc.) digitized the resultant signals in a 64-sample window that is condensed to a peak pulse value or summed pulse energy. The summed pulse energy was introduced to mitigate the quantization noise, but the experimental improvement was marginal and the simpler peak value metric therefore served as the basis for the Anger logic calculations. Peak values near the 0 and 1V digitization limits were before computing the total counts, and the Anger-determined event positions were mapped to the corresponding scintillator fibers. Multi-channel analysis was performed and displayed using a LabView program with 5-second updates of the positron event count maps.
Fig. 3. Block diagram of detection circuitry. The position sensitivity PMT and front end Anger-logic electronics are housed in a shielded box to reduce electromagnetic interference and stray light.
3. System characterization
3.1 Spatial responsivity and position mapping
In order to localize radiotracer distributions within the volume scanned by the OCT catheter, the nuclear detectors must have sufficient sensitivity and overlap between individual responses within this region. To evaluate the field of view of each detector and the entire probe, the spatial response was measured by scanning a 3 kBq Tl204 source in the lateral and longitudinal directions. The 1.0 mm diameter source was placed 1.5 mm from the probe surface and scanned with a 0.25 mm step size. The relative sensitivity of each tip was determined from the peak counts for each fiber during a calibration scan in the longitudinal direction and all subsequent nuclear counts corrected to match the lowest sensitivity detector.
Figure 4 depicts representative (a) longitudinal and (b) lateral spatial responses. The uncorrected plots demonstrate a relative sensitivity variation of less than 30%. This variation can be readily normalized without introduction of large statistical errors. The maximum counting rates were about 500 counts per minute or 3 cps/kBq for each detector. The individual responses show symmetrical characteristics with low crosstalk contamination from other fibers (< 1% longitudinal, < 4% across probe). With a longitudinal spacing of 1.75 mm, the response curves overlap at the 60–70% sensitivity points in air, which enables determination of the source locations at intermediate positions. In contrast, the wide lateral spacing required for the OCT slot limits the overlap between the detector response curves to less than 10%. Use of smaller diameter OCT fibers (see discussion) will increase the overlap to the 50% sensitivity points.
Fig. 4. Representative detector responses for (a) longitudinal and (b) lateral scans of a 1mm Tl204 source. Error bars denote standard deviations for eight 30-second measurements.
For position determination, a simple centroid algorithm requiring no calibration factors was adopted. As each positron decay event is detected by at most one scintillator tip in the probe (neglecting gamma events), localization depends upon the statistical distribution of the detection counts in the detector array. Because the probe does not have detectors in the OCT region, some assumptions are necessary to infer the source position from surrounding detectors. The centroid approach assumes a single source location and estimates the position (x̄, ȳ) from a weighted sum of the detector counts (Ci) and locations (xi, yi) for detector i:
From the spatial response results, source position localization is significantly better in the longitudinal direction compared to the lateral direction. Fig. 5(a) displays the estimated longitudinal source position for several lateral offsets between the two fiber rows using the centroid algorithm. Actual locations are presented in red. The positioning linearity is very good between the detectors, even at displaced intermediate lateral positions where the absolute counts are reduced by 50–90%. From this response, localization of the source to within 1 mm in the longitudinal direction is possible. For source positions beyond the detector locations, the estimated locations are first pulled to the nearest detector and, at greater displacements, migrate to the center as the counts fall to background levels. In the lateral direction, however, the larger spacing limits localization to the vicinity of each detector row as illustrated in Fig. 5(b). Sources at central lateral positions can be localized but these positions exhibit very low sensitivity (< 10% of the maximum), limiting the effectiveness of the current probe in these regions. Reduction in the OCT catheter diameter, as outlined in the discussion, should improve the sensitivity in the central location by up to five times.
Fig. 5. Centroid estimated position vs. true position for a scanned Tl204 beta source. (a) Scan along longitudinal direction. The rows are spaced 1.5 mm in the lateral dimension with the first and fourth rows corresponding to the two rows of scintillator detector tips. The plots demonstrate the excellent longitudinal position determination within the 3 to 9 mm range of the tips achieved for all lateral positions. (b) Scan along the lateral direction. The columns were spaced 0.8 mm apart with the first and third coinciding with opposing sets of detectors across the OCT channel. The small spatial overlap of the detector responses in this dimension currently permits only localization to either side of the OCT channel. Error bars denote standard deviations for eight 30-second measurements.
3.2 Gamma subtraction
The positron detectors are not only sensitive to positrons but also to gamma rays generated from positron annihilation in surrounding tissue. The high gamma background can therefore produce high background levels even with a positron-to-gamma sensitivity ratio of 10:1 to 20:1 typical of low-Z plastic scintillators [28]. In order to reduce the impact of gamma contamination, separate Cu-shielded fibers measure the approximate gamma background levels that are then subtracted from the detector counts [7, 8]. Scaling factors for each fiber were estimated during calibration by using a distant Cu-shielded Cs137 disk acting as a nearly uniform background source.
The effectiveness of the gamma subtraction technique was evaluated by performing spatial scans of a 1 mm Tl204 beta particle disk source between the probe and a 180 kBq Cs137 gamma source. Figure 6(a) displays the spatial response curves using the raw measured counts. Although the gamma source possesses an activity approximately 60 times higher than the beta source, the lower gamma sensitivity of the plastic tips and the larger distance from the detectors results in a background detected at 20% of the maximum beta response. Subtraction of the measured gamma contamination using the Cu-shielded fibers was effective in reducing the background levels to < 5% of the peak value (Fig. 6(b)). Figure 6(c) presents the corresponding position mapping curves for both lateral and longitudinal directions demonstrating the improvement in positioning accuracy. Similar localization linearity was observed as in the gamma-free configuration (Fig. 5(a)).
Fig. 6. Raw (a) and gamma subtracted (b) counts for a 1.5 mm diameter Tl204 source scanned in the longitudinal direction. Centroid position determination vs. true position (c) using gamma-subtracted data set and raw data sets. Error bars denote standard deviations for eight 30-second measurements.
3.3 Sensitivity (with 18F-FDG)
In preparation for the pre-clinical trials, the probe sensitivity was measured using the 18F-FDG radiotracer planned for human studies. The probe was suspended above a hemispherical bowl of 1850 kBq 18F-FDG and monitored for a period of 180 minutes at 10-minute intervals. The detectors demonstrated good linearity with an estimated minimum sensitivity of about 180 Bq with the conventional estimation of the lowest statistically significant counting rate at 3 times the background rate [29] (~ 1–2 counts per 30-second interval). This value is representative of the typical values obtained from tumors as estimated from the standardized uptake ratios from positron emission tomography (PET) detectable tumors [3, 4]. In practice, the gamma subtraction process for the background will be imperfect for high background levels and may limit the system sensitivity further
3. Co-registered localization and imaging
3.1 Localization of 18F-FDG source
The ability to localize positron radiotracer was investigated with a 10 μL, 3 mm diameter hemispherical 18F-FDG source. Figure 7 presents the measured counts and estimated source locations for longitudinal scans at two lateral offsets. The initial activity of the source was approximately 4.8 kBq and all measurements were corrected for the decay during the measurement timeframe. When the source was scanned near the row of detectors, counts were greater than 200 per 15-second interval with good linearity at intermediate locations. The overall positioning accuracy and insensitivity to lateral location parallel the results obtained with the beta particle disk sources. Near the ends of the longitudinal imaging region bounded by the probe tips, the apparent positions are approximately 1 mm from the true locations. Implications of, and potential mitigating approaches for, handling of the errors from sources outside the edges of the scanning region are considered in the discussion.
Fig. 7. Centroid-estimated source position for a 3 mm 18F-FDG hemispherical source scanned along the longitudinal probe axis beneath the scintillator probe tips and displaced 1.5 mm laterally toward the OCT center. Error bars denote standard deviations for four 15-second measurements. The individual fibers are located at 3.0, 4.75,6.5, and 8.25 mm.
3.2 Subsurface source imaging
As most ovarian cancers are intraepithelial, a tissue-simulating phantom was constructed with an embedded source to validate the effectiveness of radiotracer guidance and co-registered OCT imaging. The phantom consisted of a 3-kBq Tl204 source restricted to a 1 mm diameter using a 0.3mm thick Cu plate and central opening covered with a 0.3 mm thick mouse ear. The probe was suspended about 1 mm about the tissue. OCT scans were performed at 0.25 mm spacing for several positions within the probe field-of-view. Figures 8(b) and 8(c) show the measured detector counts, computed centroid locations, and the corresponding OCT images when the source was positioned near one end of the probe and centrally with respect to the OCT slot. The OCT images clearly localize the source near the longitudinal center between fibers 4 and 9. The positron detection similarly positions the source in this location even though the absolute sensitivity in this location is quite low due to the extended distance from the fibers (> 1 mm). The centroids of 30-second subintervals show some scatter but cluster around the full 4-minute centroid indicating the effectiveness of short measurement times.
Fig. 8. Measured counts in each probe detection fiber and corresponding sector-scan OCT images at three locations spanning 1 mm demonstrating co-registered localization and imaging.
The tissue volume visible from the OCT sector scanning extends beyond the detector tips as evidenced by the image of the fiber end in the center OCT image of Fig. 8 (marker). The co-registered localization agreement therefore demonstrates promise in correlating contrast from both modalities throughout the probe field-of-view. The centroid mapping representation of the localization information provides a simpler condensation of the positron data enabling easier comparison with the structural images. The disadvantage of this approach is the assumption of a single, concentrated uptake of radiotracer. In practice, the radiotracer uptake may be heterogeneous. For this situation, the detector count map may offer a better representation of the functional contrast for the clinician than the centroid positioning as lesion irregularities can be evaluated spatially with consideration of the relative uptake within the probe domain.
4. Discussion
Intraoperative positron and gamma probes have been investigated for detection of residual disease during surgical resection following diagnosis using conventional imaging modalities such as PET, X-Ray/CT, or MRI. In contrast, the purpose of this hybrid positron/OCT probe is to assist diagnosis of early neoplastic changes during prospective prophylactic oophorectomy (PO), which has become the standard of care for high-risk ovarian cancer patients. Most microscopic cancers or pre-malignant changes are not detectable by visual inspection and existing screening methodologies commonly diagnose ovarian cancer only in late Stages III and IV. Recent studies have indicated a higher mortality rate for pre-menopausal women undergoing PO, highlighting the need for improved methods for surgical evaluation of potential cancerous changes. A second and just as important application for this technology is the problem of detecting recurrent disease in a patient already treated for ovarian cancer. Current imaging technologies lack the resolution to determine disease status unless there is a mass at least 1–2 cm in size. The resolution of this technology should be a great improvement on current technology.
Positron probes tradeoff sensitivity and detection depth with resolution. Larger detection areas improve detection of smaller radiotracer activity levels at the expense of localization. With a pixilated detection array, this probe enables localization to dimensions commensurate with the mean-free path of positrons within the tissue. Determination of elevated radiotracer uptake to within 2 mm facilitates targeting of the high-resolution OCT images that can sample only a small fraction of the ovarian surface during surgery. One important consideration for interpretation and correlation of the positron and OCT localizations is the potential aliasing of radiotracer uptake outside the region bounded by the probe tips. As evident from the scans in Figures 5 to 7, sources just outside the probe tips will be positioned within the region due to the inherent properties of the centroid algorithm. In practice, these confounding effects can be minimized by application of minimum counting thresholds for determination of suspect tissue areas and expansion of the OCT investigative volume to account for uncertainties near boundaries of the probe field-of-view. In addition, the probe can be repositioned if uncertainties near boundaries occur.
The current probe was designed to demonstrate and evaluate the concept of radiotracer guidance and co-registered optical imaging. Efforts are currently underway to reduce the probe dimensions to fit through a standard 5 mm laparoscope accessory port. Prototypes of an OCT fiber with integrated lens and side reflector with outer dimensions less than 1 mm have shown promising performance (Fig. 9). With improved packaging, this catheter will be suitable for translational and rotational scanning within the probe. This reduction will allow decrease of the lateral gap to less than 3 mm, increasing the ability for lateral localization as well as lowering the positron sensitivity penalty for the central locations to less than 50% from the peak. In addition, with the use of prisms, the scintillating fibers can exit the probe along the longitudinal axis as opposed to orthogonally.
Fig. 9. OCT fiber with integrated angled lens: (a) Photograph and (b) Image of porcine ovary epithelium obtained with probe. The white arrows indicate ovarian follicles.
The detection sensitivity of approximately 3 cps/kBq for each pixel is similar to that reported by Yamamoto et al. [17] with a probe that also used plastic scintillators. Due to the possibility of multiple suspect regions to be interrogated intraoperatively, a radiation integration time of less than 30 seconds per location is desirable. Preliminary measurements reported herein indicate that 30-second measurement times should be achievable for radiotracer activities of about 200 Bq near the fiber detectors although the effectiveness of the gamma subtraction in a very high background could increase this time. Radiotracer uptake is a complex function of many variables including initial dose, uptake timing, tumor size, position, and local metabolism, and hormonal cycle. Future work will be focused on clinical investigation of radiotracer uptake with normal and diseased ovaries to evaluate these and other factors
5. Conclusion
A novel prototype intraoperative system combining positron detection and optical coherence tomography has been developed for ovarian cancer detection during minimally invasive surgical intervention. The optical probe features a pixilated scintillating fiber tip array for localization of 18F-FDG-radiotracer uptake to within 2 mm over a 5 × 6 mm region. An integrated high-resolution frequency domain OCT catheter, operating in a linear-plus-sector scanning mode, provides co-registered images of tissue morphology for correlation with the localized metabolic uptake determined from the positron detectors. Detailed characterization results were presented demonstrating position determination to within 1 mm longitudinally and 2 mm laterally. Dedicated fibers for gamma photon detection were shown to effectively reduce background gamma contamination from surrounding tissues and improve localization accuracy.
Testing with an 18F-FDG radiotracer indicated a minimum sensitivity of about 180 Bq and good positioning accuracy throughout the probe field-of-view. The ability to detect and image subsurface lesions was illustrated through interrogation of a beta particle source shielded by a mouse ear. The excellent co-registration of functional (positron) and morphological (OCT) features highlights the potential of the hybrid technology for detection of early or pre-malignant changes in ovarian tissue.
Acknowledgments
We thank Prof. Craig Levin of the Stanford School of Medicine for insightful consultations on positron probes and techniques and Prof. Quing Zhu acknowledges valuable discussions with Prof. Albert Sinusas at Yale University School of Medicine. This research was generously supported by the Connecticut Department of Public Health under contract DPH# 2008-0121.
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