A method for trace gas analysis is presented based on stimulated Raman scattering of the analyte in a hollow core photonic crystal fiber. A single beam pulsed laser excitation was used which realized both fiber enhanced Raman scattering and stimulated Raman scattering, that have significant advantages for improving the detection limit. A system was successfully developed based on the method and tested on hydrogen, carbon dioxide, and propene (C3H6) gases. From the results we suggest the method has great potential for analyzing trace complex gases such as volatile organic compounds, which can serve as biomarkers in human breath for lung cancer detection.
In vivo endoscopic Raman spectroscopy of human tissue using a fiber optic probe has been previously demonstrated. However, there remain several technical challenges, such as a robust control over the laser radiation dose and measurement repeatability during endoscopy. A decrease in the signal to noise was also observed due to aging of Raman probe after repeated cycles of harsh reprocessing procedures. To address these issues, we designed and tested a disposable, biocompatible, and sterile sheath for use with a fiber optic endoscopic Raman probe. The sheath effectively controls contamination of Raman probes between procedures, greatly reduces turnaround time, and slows down the aging of the Raman probes. A small optical window fitted at the sheath cap maintained the measurement distance between Raman probe end and tissue surface. To ensure that the sheath caused a minimal amount of fluorescence and Raman interference, the optical properties of materials for the sheath, optical window, and bonding agent were studied. The easy-to-use sheath can be manufactured at a moderate cost. The sheath strictly enforced a maximum permissible exposure standard of the tissue by the laser and reduced the spectral variability by 1.5 to 8.5 times within the spectral measurement range.
One challenge in facing the application of biomedical Raman spectroscopy is that the Raman signal is acquired in a dark operation room. It is inconvenient for both the operator and the patient because it is difficult for the operator to accurately and precisely locate the target in the dark environment, and the patient feels uncomfortable in such a setting. In this note, we propose a method to implement biomedical Raman measurement with an illumination source, by multiple filtering of the illumination and the collection optics. Experimental results are demonstrated on skin Raman measurement under 785-nm excitation.
KEYWORDS: Raman spectroscopy, Bronchoscopy, Charge-coupled devices, Tissues, Signal to noise ratio, Laser spectroscopy, Cancer, In vivo imaging, Spectrographs, Systems modeling
Preneoplastic lesions of the bronchial tree have a high probability of developing into malignant tumours.
Currently the best method for localizing them for further treatment is a combined white light and autofluorescence
bronchoscopy (WLB+AFB). Unfortunately the average specificity from large clinical trials for this combined detection
method is low at around 60%, which can result in many false positives. However a recent pilot study showed that adding
a point laser Raman spectroscopy (LRS) measurement improved the specificity of detecting lesions with high grade
dysplasia or carcinoma in situ to 91% with a sensitivity of 96% compared to WLB+AFB alone. Despite this success,
there is still room for much improvement. One constant need is to find better ways to measure the inherently weak
Raman emissions in vivo which will result in even better diagnostic sensitivity and specificity.
With this aim in mind a new generation Raman system was developed. The system uses the latest charge
coupled device (CCD) with low noise, and fast cool down times. A spectrometer was incorporated that was able to
measure both the low and high frequency Raman emissions with high resolution. The Raman catheter was also
redesigned to include a visible light channel to facilitate the accurate indication of the area being measured. Here the
benefits in the adjunct use of LRS to WLB + AFB are presented, and description of the new system and the
improvements it offers over the old system are shown.
Raman spectroscopy can provide information about the molecular composition of tissues, with potential to be applied as a diagnostic tool in lieu of histopathology. Our objectives are to determine if laser Raman spectra (RS) can be acquired reliably from the oral mucosa of patients, and to determine if the RS signature of normal oral mucosa is reproducible among anatomic oral sites and among subjects of different races and gender. 25 Caucasian and 26 Asian subjects are studied using RS with a signal acquisition time of 1 s at seven specified sites within the mouth. Multivariate analysis is used to determine the variability between tissue types and between races and gender. Unique spectra are defined for various sites in the mouth and are likely related to the degree of keratinization. However, spectral concordance by site is not greatly influenced by subject ethnicity or gender. We demonstrate, for the first time, the potential in-vivo application of RS for oral mucosal disease and demonstrate its specificity for particular mucosal types in the mouth. RS offers the potential to provide a diagnosis of disease using a noninvasive, convenient, sensitive technology that provides immediate results.
Our previous results from Raman spectroscopy studies on ex vivo lung tissue showed the technique had great
potential to differentiate between samples with different pathologies. In this work, a fast dispersive-type near-infrared
(NIR) Raman spectroscopy system was developed to collect real-time, noninvasive, in vivo human lung
spectra. The 785 nm excitation, and the collection of tissue emission were accomplished by using a reusable
fiber optic catheter which passed down the instrument channel of a bronchoscope. Filters in two stages blocked
laser emission other than 785 nm from reaching the tissue surface, and reduced fiber fluorescence and elastically
scattered excitation light from being passed to the spectrometer. The spectrometer itself consisted of one of
two holographic gratings with usable frequency ranges of: 700 to 2000 cm-1 and 1500 to 3400 cm-1. The
dispersed light was detected by a cooled CCD array consisting of 400 by 1340 pixels. To increase the resolution
of the system, while maximizing the throughput, a second fiber bundle, consisting of 54×100 μm diameter fibers
connected the catheter to the spectrometer. The fibers in this second bundle were spread out to form a parabolic
arc which replaced the conventional entrance slit. This geometry corrected for image aberrations, permitting
complete CCD vertical binning, thereby yielding up to a 20-fold improvement in signal-to-noise ratio. The
estimated spectral resolution of the system was 9 cm-1 for both gratings. So far we have measured spectra from
20 patients and have seen clear differences between spectra from tumor and normal tissue.
Confocal micro-Raman spectroscopy is used to probe the nuclei of normal human epidermal cells and epidermally derived cancer cells from nodular basal cell carcinomas. Clear differences are seen between the spectra. The nuclei of tumor cells appear to have different contributions from nucleic acids, histones, and proteins with an actin-like spectrum than those of normal epidermal cells. Changes in the contribution of DNA to the spectra are consistent with the staining of conventional histopathologic specimens. We also obtain spectra of the dermis, where it is found that the dermis close to tumor boundaries is not simply deficient in collagen, but shows signs of structural changes as well.
Micro-Raman spectroscopy covering a frequency range from 200 to 4000 cm-1 was used to image human skin melanocytes and keratinocytes with a spatial resolution of 0.5 μm. The cells were either cultivated on glass microscope slides or were located within thin sections of skin biopsies mounted on low fluorescence BaF2. A commercially available system was used to obtain the spectra utilizing a x100 long working distance objective with a numerical aperture of 0.8, and a cooled CCD. Both 633 and 515 nm excitations were tried, although the latter proved to be more effcient at producing Raman emission mostly due to the 1/λ4 dependence in light scattering. Fluorescence emission from the cells was surprisingly low. The excitation power at the sample was kept below about 2 mW to avoid damaging the cells; this was the limiting factor on how quickly a Raman image could be obtained. Despite this diffculty we were able to obtain Raman images with rich information about the spectroscopic and structural features within the cytoplasm and cell nuclei. Differences were observed between the Raman images of normal and malignant cells. Spectra from purified DNA, RNA, lipids, proteins and melanin were obtained and these spectra were compared with the skin cell spectra with the aim of understanding how they are distributed over a cell and how the distribution changes between different cells.
A less invasive method of reliably detecting skin cancers is
required. Raman spectroscopy is just one of several spectroscopic
methods that look promising, but are not yet sufficiently
reliable. More information is needed on how and why the Raman
spectra of cancerous skin tissue is different from its normal
counterpart. We have used confocal micro-Raman spectroscopy with a
spatial resolution of about a micron to obtain spectra of
unstained thin sections of human skin. We found that there were
clear differences in the Raman spectra between cancerous and
non-cancerous tissue both in cells and in the connective tissue.
The DNA contribution to the spectra was generally stronger in
malignant cells than normal ones. In regions of the dermis far
away from the tumor one obtains the usual collagen spectra of
normal skin, but adjacent to the tumor the spectra no longer
appeared to be those of native collagen.
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