The LED data had significantly higher SNRs in fewer averages than LD data for both the in−vitro and in−vivo recordings. Since instrument noise approached the signal size, it was important to reduce the noise sources as much as possible; but, the biggest gains in SNR were obtained by using LEDs (). In−vitro, the peak signal size was greater than the LED and photodiode noise contributions, yielding clear signals in few or no averages. The high noise of the LD necessitated slightly lower intensity illumination to stay within the dynamic range of the photodiode amplifier, and thus the signal size was smaller; however, the SNR was much smaller, requiring 100 or more averages to see the signal. Similarly in−vivo, signal size was smaller than total noise for both LED and LD trials, but SNRs for LED trials were higher because of lower noise.
At low intensities, the RMS/RF curves for all sources looked identical, rapidly decreasing with increasing radiant flux, which was expected as the relative contributions of shot noise and photodiode dark current decreased relative to other noise sources at higher intensities. When plotting RMS noise against radiant power, the laser’s sharp increase in RMS at 0.5 mW reflects noise introduced as the LD transitioned from LED mode to lasing. After the transition to lasing, the LD’s RMS remained significantly higher than LED or SLD RMS, implicating speckle as a possible disadvantage of coherent light sources. Although speckle interference patterns were not directly characterized in these experiments, the sharp increase in RMS noise concomitant with the laser’s transition from LED mode to lasing suggests that speckle arising from self interference of coherent light may have contributed significantly to the LD instrument noise. Additionally, our previous studies using CCD imaging technology (Rector et al., 2001
) use LEDs specifically because we observed a significant speckle pattern in the images obtained with LDs.
Similar to the non−coherent LED, the noise profile of the semi−coherent SLD increased linearly with increasing radiant flux, indicating the absence of a transition point in this source. Above intensities dominated by shot noise, SLD RMS increased at a faster rate than LED RMS (), suggesting that speckle noise may have also contributed significantly to the SLD noise profile or that this light source is less stable. The intensity−normalized FFTs of LD, SLD, and LD signals in−vitro () showed that the LD had frequency power two orders of magnitude greater than the LED or SLD across all frequencies, again reflecting contributions from white noise sources such as speckle.
For the in−vitro recordings, the similarity between the live nerve and dead nerve FFTs suggests that any cellular or molecular level noise contributions fell below the sensitivity of the recording method. In−vivo, two peaks present in the alive rat FFT and absent in the euthanized rat FFT corresponded to physiological noise generators: cardiac at 3.5 Hz and respiration at 1.0 Hz. Both LED and LD FFTs also had power in the low frequencies corresponding to 1/f noise; however, the low frequency density was greater in the alive rats than in euthanized rats, suggesting that other low frequency physiological sources may have contributed to the low frequencies. For example, in−vivo signals may also contain Mayer waves (f = 0.1 – 0.4 Hz, Mayhew et al., 1996
), spontaneous oscillations of arterial pressure tightly coupled with the oscillations of efferent sympathetic nervous system activity (for review, see Julien, 2006
). However, in optical recordings, 1/f noise dominates at these low frequencies, and we attempted to separate the effects of Mayer waves from the 1/f contribution by subtracting live and dead animal recordings for frequencies between 0 and 0.4 Hz ().
As LD intensity increased, so did the noise. Since the input range of the photodiode system is limited, we could only increase the laser diode intensity until the variability, perhaps due to the speckle noise, began to saturate the dynamic range of the photodiode system. Ideally, the brighter the light source, the bigger the SNR of the optical signal by overcoming shot noise limitations; however, the added noise of the LD counteracted gains over shot noise achieved with brighter illumination.
Since physiological noise sources were low in−vitro, higher SNRs were observed due to lower overall RMS noise. The LED SNR profile peaked at 150 averages, but then decreased over the next 850 averages because nerve health and corresponding signal amplitude decayed over time. Laser diode noise failed to increase with the square root of the number of trials like LED and SLD SNRs because below 200 averages, the signal size was significantly below LD noise levels, causing the SNRs in the 1–200 trial range to appear variable. The reduced light source noise of LEDs enabled signal recordings with higher SNRs in fewer averages, an advantage for limited−life, in−vitro preparations.
The high variability in the optical response amplitude across different rats likely resulted from variables such as the position of the photodiode over the whisker barrels and anesthesia depth. Like the in−vitro experiments, the higher SNRs of hemodynamic signals with the LED over the LD indicate that LEDs are a better light source for non−coherent neural recording applications. In spite of long integration times typically used to reduce speckle noise for the slow hemodynamic signals seen here, LEDs maintain a significant advantage, perhaps due to reduction of noise at low frequencies.
Laser diodes are a commonly used illumination source for in−vivo optical neural imaging and recording in freely moving chronic studies because of their high intensity illumination capabilities; however, we observed that the noise introduced by LDs counteracted benefits of high brightness when compared with low−noise LEDs. LED recordings produced roughly a 10−fold increase in SNRs in−vitro () and 2− to 3−fold increase in−vivo () over LD recordings. These gains are significant; however, methods must be pursued to further increase the contrast between intrinsic signals and noise, making single pass measurements more practical. Such efforts may include the development of brighter, more stable non−coherent sources and the application of birefringence recording and imaging in−vivo. Additionally, the LD intensity was limited by laser noise, with peak−to−peak fluctuations that saturated the dynamic range of the amplifiers, both in−vitro and in−vivo. LD SNRs could be improved by reducing the amplifier gain and powering the LD with higher light output. The will efecitvely reduce the proportional contribution of shot noise since it appears from that speckle remains constant as LD intensity increases. However, even when the laser was driven in its optimal range, as shown at the rightmost points in , the LD was significantly noisier than the LED for both raw () and intensity normalized () RMS noise, possibly due to speckle.
We have previously investigated methods to partially decohere the light (also discussed in McKechnie, 1984
). However, these efforts usually made speckle worse by breaking the coherent beam into more components, or the method both reduced intensity and increased complexity to the point that made lasers impractical for in−vivo measurements, especially in freely moving animals, further supporting the use of LEDs. Running a coherent source through a long, multimode fiber has been a useful method for reducing speckle noise in optical coherence tomography (Kim et al., 2005
). Coupling through a short, multimode fiber propagated coherent light and speckle patterns, and coupling through a single mode fiber cut the intensity to a point which made in−vivo recording impractical. Once the coherent light from the single mode fiber interacted with tissue, the interference patterns began to appear again.
Since light is diffuse in tissue due to scattering, a large single channel detector with a 1 cm or greater diameter might prevent speckle patterns from crossing the edge of the detector and to prevent aperture effects. We used a large area detector for the lobster experiments, but this was impractical for in−vivo, freely moving animal studies. Additionally, speckle interference patterns vary both laterally (xy) and axially (z) in tissue with respect to the detector. Using a large single−channel detector could reduce consequences of detection aperture in x and y, but not in z. Thus, the most effective way to rid signals of speckle noise in scattering tissue is to illuminate with noncoherent light.
Independent of detection aperture effects, constructive interference of coherent light produces bright spots that can saturate points in the middle of the detector, resulting in an output that does not accurately reflect average intensity across the detector. Thus, when these saturating speckles move around the detector, the output will still fluctuate greatly because the photodetector does not accurately measure the total light intensity across the area. If a tightly focused beam of coherent light is not required for an optical measurement (i.e., near infrared diffuse optical tomography and scattered light imaging), non−coherent light is preferred because coherent light speckle may introduce excessive noise. Since low−coherence SLDs offer a bright and narrow beam alternative to highly coherent LDs, we included this light source in our analysis.
While mercury−, xenon−, and mercury−xenon arc lamps are commonly used for fluorescence and voltage sensitive dye applications in microscopy (Zochowski et al., 2000
), for this study we selected LEDs over halogen light for a non−coherent source for in−vivo measurements because LEDs are more convenient, just as bright at narrow wavelengths (Foust et al., 2005
), and more stable at low frequencies (Rumyantsev et al., 2004
) than halogen sources. Some studies have also shown that LEDs have advantages over halogen sources for intrinsic and voltage sensitive dye recordings (Salzberg et al., 2005
; Nishimura et al., 2006
). LED technology improves each year and has created low power, efficient devices that may eventually exceed intensities possible by halogen and arc lamp sources for narrow wavelengths.
While lasers have been practical for detecting slow hemodynamic changes (i.e., Villringer and Chance, 1997
; Boas, 2002
), low noise sources such as LEDs are critical in advancing toward single pass measurements of rapid intrinsic optical signals in−vivo. For measurement of slow hemodynamic changes, long integration times (~100 ms) may average out the speckle noise generated in coherent laser beams. However, in measurements of rapid signals with short integration times (10 ms or faster), the speckle noise may swallow the small, transient changes or even the more robust hemodynamic changes recorded in single passes. A close look at the literature over the past year shows many systems use laser diodes for in−vivo chronic measurements of hemodynamic changes. Currently available commercial devices also use laser diodes (e.g., Hamamatsu NIRO−200, Hitachi ETG−4000). Several investigators use halogen sources (e.g., Shtoyerman et al., 2000
; Brett−Green et al., 2001
; Roe, 2007
) for fixed imaging of acute preparations, or restrained preparations with windows. Only recently have some investigators published work with LEDs (e.g., Chen−Bee et al., 2007
; Zeff et al., 2007
). Recent studies utilizing LEDs in−vivo have achieved some of the most robust signals ever observed, presumably due to the advantages of LEDs highlighted by the present study and by Salzberg et al. (2005)
For freely moving human and animal studies (e.g., Chance et al., 1998
; Gratton and Fabiani, 2003
), lasers have been popular because LDs with output powers in the FDA approved range (5–30 mW) can be modulated at high frequencies for phase measurements. In the current study, all three light sources could produce about the same output intensity (~4mW, ). Indeed, there are lasers that are brighter than our LEDs, however, LEDs also share rapid modulation capabilities and are becoming as bright (10−60 mW) as class IIIb lasers in the near−infrared wavelengths, which is encouraging for in−vivo work as light in this frequency range penetrates skin and skull and is scattered by neural tissue (Eggert and Blazek, 1987
). For comparison purposes, we used the brightest LED available at the wavelength of our laser diode (~660 nm), as this wavelength was optimal for detecting the hemodynamic changes in−vivo. Brighter near infrared LEDs are currently available, and our preliminary work with these LEDs indicates that they maintain low noise characteristics.
There are several factors that recommend LEDs as practical, inexpensive, and low−noise alternatives to coherent sources for neural recording applications. The sharp increase in RMS noise at lasing threshold suggests that speckle noise contributed significantly to the LD’s lower SNRs. LEDs exhibit greater, low−frequency stability than lasers and halogen sources (Rumyantsev et al., 2004
; Salzberg et al., 2005
). We tried using both voltage and current sources to drive the LEDs and LDs with no differences in the RMS noise levels. A current regulated power source would minimize very slow variations in intensity due to temperature and drift, but these effects were not significant in our studies. LEDs do not normally require temperature regulation, and can be driven with low power voltage sources, although current sources and temperature regulation may be preferred for applications requiring long−term stability. Many techniques use fiber coupled light from a variety of different sources including halogen, xenon and lasers for delivering photons directly to the animal. However, fiber coupling is bulky and restrictive to animal movement, especially if the animal is as small as a rat or a mouse. Additionally, movement of the fiber results in additional noise in the illumination intensity due to changing photon paths through the fiber and changes in the coupling of the fiber to the animal. In contrast, LEDs are inexpensive, easy to use, and can be easily carried by animal and human subjects.