Limits...
Breaking the spatial resolution barrier via iterative sound-light interaction in deep tissue microscopy.

Si K, Fiolka R, Cui M - Sci Rep (2012)

Bottom Line: Random scattering causes the ballistic focus, which is conventionally used for image formation, to decay exponentially with depth.Optical imaging beyond the ballistic regime has been demonstrated by hybrid techniques that combine light with the deeper penetration capability of sound waves.This development opens up practical high resolution fluorescence imaging in deep tissues.

View Article: PubMed Central - PubMed

Affiliation: Howard Hughes Medical Institute, Janelia Farm Research Campus, 19700 Helix Drive, Ashburn, Virginia 20147, USA.

ABSTRACT
Optical microscopy has so far been restricted to superficial layers, leaving many important biological questions unanswered. Random scattering causes the ballistic focus, which is conventionally used for image formation, to decay exponentially with depth. Optical imaging beyond the ballistic regime has been demonstrated by hybrid techniques that combine light with the deeper penetration capability of sound waves. Deep inside highly scattering media, the sound focus dimensions restrict the imaging resolutions. Here we show that by iteratively focusing light into an ultrasound focus via phase conjugation, we can fundamentally overcome this resolution barrier in deep tissues and at the same time increase the focus to background ratio. We demonstrate fluorescence microscopy beyond the ballistic regime of light with a threefold improved resolution and a fivefold increase in contrast. This development opens up practical high resolution fluorescence imaging in deep tissues.

No MeSH data available.


Related in: MedlinePlus

(a) Lateral PSF measurement through 2 mm thick tissue phantoms (μs  = 7.63 /mm, g factor = 0.9013) for iterations 1, 3, 5, 7, and 9.To normalize the peak intensity, the PSF data sets were multiplied by 6.5, 2, 1.5, and 1.5 for iteration 1, 3, 5, and 7, respectively. (b) Axial PSF measurements for iterations 1, 5, and 9. The PSF data for iteration 1 and 5 was multiplied by 6.5 and 1.5, respectively. (c–e) Gaussian fitting of the measured PSF. (f) Fitted transverse FWHM and simulation (mean values and standard deviation). (g) Fitted axial FWHM and simulation (mean values and standard deviation). (h) Measured focus to background ratio and simulation (mean values and standard deviation). (i) Measured ultrasound modulated light power and simulation (mean values and standard deviation). Scalebar: 10 microns. Colorbar in arbitrary units.
© Copyright Policy - open-access
Related In: Results  -  Collection

License
getmorefigures.php?uid=PMC3475990&req=5

f2: (a) Lateral PSF measurement through 2 mm thick tissue phantoms (μs = 7.63 /mm, g factor = 0.9013) for iterations 1, 3, 5, 7, and 9.To normalize the peak intensity, the PSF data sets were multiplied by 6.5, 2, 1.5, and 1.5 for iteration 1, 3, 5, and 7, respectively. (b) Axial PSF measurements for iterations 1, 5, and 9. The PSF data for iteration 1 and 5 was multiplied by 6.5 and 1.5, respectively. (c–e) Gaussian fitting of the measured PSF. (f) Fitted transverse FWHM and simulation (mean values and standard deviation). (g) Fitted axial FWHM and simulation (mean values and standard deviation). (h) Measured focus to background ratio and simulation (mean values and standard deviation). (i) Measured ultrasound modulated light power and simulation (mean values and standard deviation). Scalebar: 10 microns. Colorbar in arbitrary units.

Mentions: Figure 2 a shows the lateral PSF for DOPC iteration 1, 3, 5, 7, and 9. Figure 2 b shows the axial PSF for iteration 1, 5, and 9. To determine the full width half maximum (FWHM), Gaussian fitting through cross-sections of each PSF was applied (Fig. 2 c–e). For iteration 1, when DOPC is applied for the first time, the mean FWHM of the PSF amounts to 35.7, 39.0 and 142 microns in the y, z and x (axial) direction, respectively (Fig. 2 f–g). After nine DOPC iterations, the FWHM was reduced to 11.2, 12.8, and 60.3 microns in the y, z, and x directions. The FBR is increased by a factor of ~5 over nine iterations and appears to grow almost linearly with N (Fig. 2 h). Besides the FBR, the total sound modulated light power increases as well (Fig. 2 i), however not linearly with N. We have simulated the iterative DOPC process (see Supplementary discussion) and the results are generally in good agreement with the experiments (Fig. 2 f–i). In addition, we also performed PSF measurements through 1.2 mm thick fixed rat brain tissue, as shown in Supplementary Fig. 1.


Breaking the spatial resolution barrier via iterative sound-light interaction in deep tissue microscopy.

Si K, Fiolka R, Cui M - Sci Rep (2012)

(a) Lateral PSF measurement through 2 mm thick tissue phantoms (μs  = 7.63 /mm, g factor = 0.9013) for iterations 1, 3, 5, 7, and 9.To normalize the peak intensity, the PSF data sets were multiplied by 6.5, 2, 1.5, and 1.5 for iteration 1, 3, 5, and 7, respectively. (b) Axial PSF measurements for iterations 1, 5, and 9. The PSF data for iteration 1 and 5 was multiplied by 6.5 and 1.5, respectively. (c–e) Gaussian fitting of the measured PSF. (f) Fitted transverse FWHM and simulation (mean values and standard deviation). (g) Fitted axial FWHM and simulation (mean values and standard deviation). (h) Measured focus to background ratio and simulation (mean values and standard deviation). (i) Measured ultrasound modulated light power and simulation (mean values and standard deviation). Scalebar: 10 microns. Colorbar in arbitrary units.
© Copyright Policy - open-access
Related In: Results  -  Collection

License
Show All Figures
getmorefigures.php?uid=PMC3475990&req=5

f2: (a) Lateral PSF measurement through 2 mm thick tissue phantoms (μs = 7.63 /mm, g factor = 0.9013) for iterations 1, 3, 5, 7, and 9.To normalize the peak intensity, the PSF data sets were multiplied by 6.5, 2, 1.5, and 1.5 for iteration 1, 3, 5, and 7, respectively. (b) Axial PSF measurements for iterations 1, 5, and 9. The PSF data for iteration 1 and 5 was multiplied by 6.5 and 1.5, respectively. (c–e) Gaussian fitting of the measured PSF. (f) Fitted transverse FWHM and simulation (mean values and standard deviation). (g) Fitted axial FWHM and simulation (mean values and standard deviation). (h) Measured focus to background ratio and simulation (mean values and standard deviation). (i) Measured ultrasound modulated light power and simulation (mean values and standard deviation). Scalebar: 10 microns. Colorbar in arbitrary units.
Mentions: Figure 2 a shows the lateral PSF for DOPC iteration 1, 3, 5, 7, and 9. Figure 2 b shows the axial PSF for iteration 1, 5, and 9. To determine the full width half maximum (FWHM), Gaussian fitting through cross-sections of each PSF was applied (Fig. 2 c–e). For iteration 1, when DOPC is applied for the first time, the mean FWHM of the PSF amounts to 35.7, 39.0 and 142 microns in the y, z and x (axial) direction, respectively (Fig. 2 f–g). After nine DOPC iterations, the FWHM was reduced to 11.2, 12.8, and 60.3 microns in the y, z, and x directions. The FBR is increased by a factor of ~5 over nine iterations and appears to grow almost linearly with N (Fig. 2 h). Besides the FBR, the total sound modulated light power increases as well (Fig. 2 i), however not linearly with N. We have simulated the iterative DOPC process (see Supplementary discussion) and the results are generally in good agreement with the experiments (Fig. 2 f–i). In addition, we also performed PSF measurements through 1.2 mm thick fixed rat brain tissue, as shown in Supplementary Fig. 1.

Bottom Line: Random scattering causes the ballistic focus, which is conventionally used for image formation, to decay exponentially with depth.Optical imaging beyond the ballistic regime has been demonstrated by hybrid techniques that combine light with the deeper penetration capability of sound waves.This development opens up practical high resolution fluorescence imaging in deep tissues.

View Article: PubMed Central - PubMed

Affiliation: Howard Hughes Medical Institute, Janelia Farm Research Campus, 19700 Helix Drive, Ashburn, Virginia 20147, USA.

ABSTRACT
Optical microscopy has so far been restricted to superficial layers, leaving many important biological questions unanswered. Random scattering causes the ballistic focus, which is conventionally used for image formation, to decay exponentially with depth. Optical imaging beyond the ballistic regime has been demonstrated by hybrid techniques that combine light with the deeper penetration capability of sound waves. Deep inside highly scattering media, the sound focus dimensions restrict the imaging resolutions. Here we show that by iteratively focusing light into an ultrasound focus via phase conjugation, we can fundamentally overcome this resolution barrier in deep tissues and at the same time increase the focus to background ratio. We demonstrate fluorescence microscopy beyond the ballistic regime of light with a threefold improved resolution and a fivefold increase in contrast. This development opens up practical high resolution fluorescence imaging in deep tissues.

No MeSH data available.


Related in: MedlinePlus