Biomedical Engineering Reference
In-Depth Information
1.0
WT simulated
WT experimental
0.9
0.8
0.7
0.6
0.5
Oim experimental
0.4
Oim simulated
0.3
0
10
20
30
40
50
60
Depth (microns)
FIgurE 6.8
Comparison of the experimental forward SHG attenuation data with Monte Carlo simulations (with
associated standard errors) based on the bulk optical parameters at both the fundamental and SHG wavelengths
(Table 6.1). The creation directionality was taken from Figure 6.7b, and relative SHG conversion efficiency of 2.54-
fold larger for the WT was used. As absolute magnitude of the SHG intensity from the oim is smaller than that of the
WT, the data are normalized to their respective maximum and also to the maximum in each series to account for
local variability in the tissues. Chi-squared tests for both the WT and oim indicate that the respective experimental
and simulated results are not significantly different. (Reprinted from
Biophys. J.
94, Lacomb, R., O. Nadiarnykh,
and P. J. Campagnola, Quantitative SHG imaging of the diseased state osteogenesis imperfecta: Experiment and
simulation, 4504-4514, Copyright (2008), with permission from Elsevier.)
and dominates the observed response. The measured intensity is further determined by the extent of the
remaining tissue the photons must travel through to be collected (secondary filter effects); however, in all
tissues, we have examined this is a small effect in this wavelength range and tissue thickness.
We must also consider that the relative SHG intensity from the oim skin (using the same laser excita-
tion power) is weaker than that from the WT. Thus, the normalized SHG intensity from the oim will
decay faster relative to the WT due to fewer initially generated photons at subsequent depths. A similar
mechanism was proposed by Welch et al. [41] for fluorescence measurements in tissue, where they intro-
duced the idea of “weighted photons” that accounted for local absorption coefficients and fluorescence
quantum yields. We draw upon this idea to compare the SHG signal propagation in these different
SHG-producing tissues. Rather than an absorption coefficient, the SHG intensity is determined by the
second-order nonlinear susceptibility χ
(2)
. While we do not determine absolute χ
(2)
coefficients, as the
apparent efficiency will be convolved with scattering for tissues of greater thickness than 1 MFP, we
can estimate the relative conversion efficiencies for the WT and oim tissues based on SHG intensity
measurements. We cannot measure these values directly in skin as it is not possible to slice specimens of
insufficient thickness (<30 microns) such that the initial SHG intensities can be measured in the absence
of scattering. As an alternative, we performed these measurements in thin oim and WT bone slices
(6 microns), where multiple scattering will be insignificant. This approach assumes similar changes
in the collagen between skin and bone in the diseased state. Additionally, bone cryosections are fairly
uniform, whereas analogous sections of skin can display substantial cutting nonuniformities. These
measurements reveal that the WT was ~2.54 ± 0.22 (
p
= 0.04)-fold brighter than the oim bone [12]. We
previously reported that the collagen concentration in these tissues was similar (based on quantitative
Sirius Red staining); thus, the observed intensity differences can be ascribed to the difference in χ
(2)
. As
χ
(2)
is the spatially averaged macroscopic analog of the molecular hyperpolarizability, β, it is expected
to have a lower value in the more disordered tissue, even if β is the same between the tissues. It is likely,
however, that the β will also be different due to changes in helical structure in the diseased tissue.
