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Optics Express

Optics Express

  • Editor: Michael Duncan
  • Vol. 13, Iss. 12 — Jun. 13, 2005
  • pp: 4611–4628
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Form-biattenuance in fibrous tissues measured with polarization-sensitive optical coherence tomography (PS-OCT)

Nate J. Kemp, Haitham N. Zaatari, Jesung Park, H. Grady Rylander III, and Thomas E. Milner  »View Author Affiliations


Optics Express, Vol. 13, Issue 12, pp. 4611-4628 (2005)
http://dx.doi.org/10.1364/OPEX.13.004611


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Abstract

Form-biattenuance (Δχ) in biological tissue arises from anisotropic light scattering by regularly oriented cylindrical fibers and results in a differential attenuation (diattenuation) of light amplitudes polarized parallel and perpendicular to the fiber axis (eigenpolarizations). Form-biattenuance is complimentary to form-birefringence (Δn) which results in a differential delay (phase retardation) between eigenpolarizations. We justify the terminology and motivate the theoretical basis for form-biattenuance in depth-resolved polarimetry. A technique to noninvasively and accurately quantify form-biattenuance which employs a polarization-sensitive optical coherence tomography (PS-OCT) instrument in combination with an enhanced sensitivity algorithm is demonstrated on ex vivo rat tail tendon (mean Δχ=5.3·10-4, N=111), rat Achilles tendon (Δχ=1.3·10-4, N=45), chicken drumstick tendon (Δχ=2.1·10-4, N=57), and in vivo primate retinal nerve fiber layer (Δχ=0.18·10-4, N=6). A physical model is formulated to calculate the contributions of Δχ and Δn to polarimetric transformations in anisotropic media.

© 2005 Optical Society of America

1. Introduction

Polarization-sensitive optical coherence tomography (PS-OCT) incorporates polarimetric sensitivity with the ranging capability of optical coherence tomography (OCT) to characterize the depth-resolved polarization state of light backscattered from turbid media such as biological tissue. Using requisite processing methods, diagnostically-relevant polarization properties such as tissue form-birefringence and optic axis orientation can be decoded from the detected transformation in light polarization state versus propagation depth [1

1. C. K. Hitzenberger, E. Gotzinger, M. Sticker, M. Pircher, and A. F. Fercher, “Measurement and imaging of birefringence and optic axis orientation by phase resolved polarization sensitive optical coherence tomography,” Opt. Express 9, 780–790 (2001), http://www.opticsexpress.org/abstract.cfm?URI=OPEX-9-13-780. [CrossRef] [PubMed]

4

4. S. Jiao and L. V. Wang, “Jones-matrix imaging of biological tissues with quadruple-channel optical coherence tomography,” J. Biomed. Opt. 7, 350–358 (2002). [CrossRef] [PubMed]

]. Accurate quantification of tissue polarization properties in thin specimens exhibiting weak form-birefringence necessitates an approach which minimizes the effects of speckle noise and exploits the known behavior of polarization transformations in a noise-free model [5

5. N. J. Kemp, J. Park, H. N. Zaatari, H. G. Rylander, and T. E. Milner, “High sensitivity determination of birefringence in turbid media using enhanced polarization-sensitive OCT,” J. Opt. Soc. Am. A 22, 552–560 (2005). [CrossRef]

].

In the eigenpolarization coordinate frame, the polarization-transforming properties of a non-depolarizing, homogeneous optical medium such as anisotropic fibrous tissue are described by the Jones matrix

J=[exp((Δχ+iΔn)πΔzλ0)00exp((ΔχiΔn)πΔzλ0)]
=[ξ1exp(iarg(ξ1))00ξ2exp(iarg(ξ2))],
(1)

where ξ 1 and ξ 2 are the complex eigenvalues representing changes in amplitude and phase for orthogonal eigenpolarization states with free-space wavelength λ0 propagating a distance Δz through the medium. Attenuation common to both eigenpolarizations does not affect the light polarization state and is neglected here.

The phase retardation (δ, expressed in radians) between eigenpolarization states after propagation through the medium is the difference between the arguments of the eigenvalues, δ=arg(ξ 1)-arg(ξ 2), which allows simplification of the Jones matrix to

J=[ξ1exp(iδ2)00ξ2exp(iδ2)].
(2)

Chipman [9

9. R. A. Chipman, “Polarization analysis of optical systems,” Opt. Eng. 28, 90–99 (1989).

,10

10. S.-Y. Lu and R. A. Chipman, “Interpretation of Mueller matrices based on polar decomposition,” J. Opt. Soc. Am. A 13, 1106–1113 (1996). [CrossRef]

] introduced the polarimetric parameter diattenuation (D) given quantitatively by

D=T1T2T1+T2=ξ12ξ22ξ12+ξ220D1,
(3)

where T 1 and T 2 are the intensity transmittances for the two orthogonal eigenpolarizations and the attenuation can be a consequence of either anisotropic absorption or anisotropic scattering of light out of the detected field.

Birefringencen) is the phenomenon responsible for phase retardation (δ) of light propagating a distance Δz in an anisotropic element and is given by

Δn=λ02πδΔz=nsnf,
(4)

where ns and nf are the real-valued refractive indices experienced by the slow and fast eigenpolarizations, respectively. Note form-birefringence (Δn) is proportional to and given experimentally by the phase retardation-per-unit-depth (δz).

Δχ=χsχf,
(5)

where χs and χf are attenuation coefficients of the slow and fast eigenpolarizations. For absorbing (dichroic) media, χs and χf are simply imaginary-valued refractive indices.

For depth-resolved polarimetry in scattering media (i.e. PS-OCT) [14

14. J. F. de Boer, T. E. Milner, and J. S. Nelson, “Determination of the depth-resolved Stokes parameters of light backscattered from turbid media using Polarization Sensitive Optical Coherence Tomography,” Opt. Lett. 24, 300–302 (1999). [CrossRef]

], expression of an element’s form-biattenuance on a per-unit-depth basis is desirable both theoretically and experimentally. We define the relative-attenuation (ε) experienced by light propagating to a depth Δz in an anisotropic element as

ε=2πλ0ΔzΔχ,
(6)

and form-biattenuance (Δχ) can now be meaningfully expressed on a relative-attenuation-per-unit- depth basis (εz). Relative-attenuation (ε) is the complimentary term to phase retardation [δ, Eq. (4)], just as biattenuance (Δχ) is the complementary term to birefringence (Δn).

The Jones matrix of the anisotropic medium from Eqs. (1) and (2) becomes

J=[exp(ε+iδ2)00exp(εiδ2)],
(7)

D=eεeεeε+eε=tanh(ε).
(8)

We note that for small relative-attenuation, a small-angle approximation is valid and Dε.

Phase retardation (δ) is in the argument of an exponential [Eq. (7)] and therefore has units of radians, but is also commonly expressed in units of degrees (180·δ/π), fractions of waves (δ/2π), or length (λ0 ·δ/2π). Similarly, relative-attenuation (ε) is in the argument of an exponential [Eq. (7)] and has units of radians. Regrettably, expression of relative-attenuation in units of radians (or degrees, fractions of waves, or length) is less intuitive than for phase retardation.

Several investigators have reported PS-OCT measurements of diattenuation. Park et al [16

16. B. H. Park, M. C. Pierce, B. Cense, and J. F. de Boer, “Jones matrix analysis for a polarization-sensitive optical coherence tomography system using fiber-optic components,” Opt. Lett. 29, 2512–2514 (2004). [CrossRef] [PubMed]

] recently reported an approximate single-pass “diattenuation-per-unit-depth” in chicken tendon of Dz=0.39/mm using a PS-OCT instrument with 1310 nm light. For their sample thickness (Δz=0.4 mm), the small-angle approximation is valid and εz=0.39 rad/mm (Δχ=0.82·10-4). Unfortunately, reporting Dz=0.39/mm mistakenly implies that a 3-mm-thick specimen of the same tendon would have diattenuation D=1.17 [0≤D≤1, Eq. (3)]. Jiao and Wang [4

4. S. Jiao and L. V. Wang, “Jones-matrix imaging of biological tissues with quadruple-channel optical coherence tomography,” J. Biomed. Opt. 7, 350–358 (2002). [CrossRef] [PubMed]

] reported single-pass Dz=0.13/mm for a sample with thickness Δz=0.3 mm and Todorovic et al [15

15. M. Todorovic, S. Jiao, L. V. Wang, and G. Stoica, “Determination of local polarization properties of biological samples in the presence of diattenuation by use of Mueller optical coherence tomography,” Opt. Lett. 29, 2402–2404 (2004). [CrossRef] [PubMed]

] reported single-pass D=0.1 for a sample with thickness Δz=1 mm. Again, the small-angle approximation is valid for these cases (εz=0.13 rad/mm or Δχ=0.17·10-4 ; εz=0.1 rad/mm or Δχ=0.13·10-4), but for specimens with higher diattenuation, Dzεz [see Eq. (8)]. Expression of a specimen’s polarization-dependent attenuation in terms of form-biattenuance (Δχ) or corresponding relative-attenuation-per-unit-depth (εz) overcomes all ambiguity associated with the nonlinear diattenuation-thickness relationship and no approximations are necessary in depth-resolved polarimetry.

We present a novel technique for measuring form-biattenuance (Δχ) in highly scattering biological media. In Section 2, we give the trajectory of noise-free model depth-resolved normalized Stokes vector arcs on the Poincaré sphere for light propagation through media exhibiting both form-birefringence and form-biattenuance and give an expression for the polarimetric signal-to-noise ratio (PSNR). Section 3 describes a PS-OCT instrument and multistate nonlinear fitting algorithm used to estimate values of form-biattenuance (Δχ) given in Section 4 for a variety of specimens. In Sections 5 and 6, we discuss the expected uncertainty in measurements of Δχ, present a rudimentary physical model of form-biattenuance based on structural properties, compare measured values given here with previously reported results, and conclude with a discussion of this work’s relevance in biomedical diagnostics.

2. Theory

The continuous transformation of perfectly polarized light propagating through a scattering medium with Jones matrix given in Eq. (7) is geometrically represented by normalized Stokes parameters vs. depth [Q(z), U(z), V(z)] which trace a noise-free model polarization arc [P(z)] on the Poincaré sphere surface (Fig. 1). Trajectory of the P(z) arc is governed by a vector differential equation [17

17. J. Park, N. J. Kemp, H. N. Zaatari, H. G. Rylander III, and T. E. Milner, Dept. of Biomedical Engineering, University of Texas, 1 University Station, #C0800, Austin, TX 78712 are preparing a manuscript to be called “Differential geometry of the normalized Stokes vector trajectories in anisotropic media.”

],

dP(z)dz+(P(z)×βre)+P(z)×(P(z)×βim)=0,
(9)

where the complex differential wavevector β is a property of the medium and is defined as

β=βre+iβim=(βre+iβim)β̂.
(10)

βre and βim are real and imaginary parts of the complex differential wavenumber β,

β=βre+iβim=2πλ0(Δn+iΔχ),
(11)

and eigen-axis β^ is a unit-vector on the Poincaré sphere representing the fast eigenpolarization state which propagates through the medium without transformation.

The vector differential equation [Eq. (9)] indicates that differential movement of the P(z) arc is orthogonal to P(z) for all z. In fibrous media, form-birefringence (Δn) and form-biattenuance (Δχ) arise from the same anisotropic scattering structures so that β re and βim share a common eigen-axis (β^) and differential movement of P(z) on the Poincaré sphere due to form-birefringence P×βredz) is orthogonal to movement due to form-biattenuance [P×(P×βim)dz]. At each depth z in the medium, the instantaneous trajectory of P(z) is resolved into two orthogonal planes: 1) a plane [Π1(z)] with normal β^ which encompasses differential movement of P(z) due to Δn, and 2) a plane [Π2(z)] containing both P(z) and β^ which encompasses differential movement of P(z) due to Δχ.

Cumulative movement of P(z) on the Poincaré sphere represents the entire transformation experienced by polarized light propagating into and back from the medium (double-pass) and can be found by direct integration of Eq. (9) in the mutually orthogonal planes Π1(z) and Π2(z). The cumulative angle between Π2(0) and Π2z) is given by

2δ=2βreΔz,
(12)

and therefore 2δ is the total double-pass phase retardation of the specimen. Separation-angle [γ(z)] is defined as the angle between P(z) and β^ and moves within Π2(z) according to

γ(z)=2tan1[tan(γ(0)2)exp(2βimz)]0γ<π,
(13)

where γ(0)=cos-1[β̂·P(0)] is the initial separation-angle between β^ and the incident polarization state [P(0)]. At γz), the total double-pass relative-attenuation (2ε) is given by

2ε=2βimΔz.
(14)

Polarimetric signal-to-noise ratio (PSNR) characterizes the ability of PS-OCT to estimate δ and ε in the presence of polarimetric speckle noise with standard deviation σ speckle (Eq. 7 of Ref [5

5. N. J. Kemp, J. Park, H. N. Zaatari, H. G. Rylander, and T. E. Milner, “High sensitivity determination of birefringence in turbid media using enhanced polarization-sensitive OCT,” J. Opt. Soc. Am. A 22, 552–560 (2005). [CrossRef]

]),

PSNR=larcσspeckle,
(15)

where larc is the arc length of the noise-free model polarization arc [P(z)]. Differential arc length (dlarc) is given by

dlarc=2(βre2+βim2)12sin[γ(z)]dz,
(16)

which is integrated to provide arc length of P(z),

larc=[1+(δε)2]12[γ(0)γ(Δz)].
(17)
Fig. 1. Noise-free model polarization arc [P(z), black] and eigen-axis (β^, green) on the Poincaré sphere (left) and corresponding normalized Stokes parameters [Q(z), U(z), V(z)] vs. depth (right). Polarizations at the front [P(0)] and rear [Pz)] specimen surfaces are represented by red and blue dots respectively. (a) Pure form-birefringence causes rotation of P(z) around β^ in plane Π1 which is normal to β^. (b) Pure form-biattenuance causes translation of P(z) toward β^ in plane Π2. (c) Combined birefringence and biattenuance cause P(z) to spiral toward β^ and orthogonal planes Π1 and Π2 are therefore functions of depth [Π1(z) and Π2(z)]. Movies showing 3D nature of Poincaré sphere: a (1.21 MB); b (1.17 MB); c (1.21 MB).

When relative-attenuation is small (ε<0.1), l arc can be approximated to the first order,

larc2(δ2+ε2)12sin[γ(0)].
(18)

3. Methods

3.1 Instrumentation

Our PS-OCT instrumentation has been previously reported in detail [5

5. N. J. Kemp, J. Park, H. N. Zaatari, H. G. Rylander, and T. E. Milner, “High sensitivity determination of birefringence in turbid media using enhanced polarization-sensitive OCT,” J. Opt. Soc. Am. A 22, 552–560 (2005). [CrossRef]

] so we provide only a brief summary here. The source is a mode-locked Ti:Al2O3 laser (λ0=830 nm and ΔλFWHM=55 nm). Source light is linearly polarized at 45° to insure that incident horizontal and vertical electric field amplitudes are equal. A 50/50 non-polarizing beam splitter divides the source beam into reference and sample paths. Longitudinal scanning (1 mm deep in air, 30 Hz) is provided by a corner-cube retroreflector mounted to a loudspeaker diaphragm in the reference path. Sample path optics include a liquid-crystal variable retarder (LCVR) oriented with its fast axis horizontally (0°), x- and y- lateral scanning galvanometers, and scanning optics. For tendon specimens, an achromatic lens (f=25 mm) focuses light onto the specimen (power=3 mW). For retinal specimens, the intact cornea focuses light onto the retinal nerve fiber layer (RNFL) (power=1.8 mW) [18

18. H. G. Rylander, N. J. Kemp, J. Park, H. N. Zaatari, and T. E. Milner, “Birefringence of the primate retinal nerve fiber layer,” Exp. Eye Res. In Press, (2005).

]. A polarizing beam splitter divides horizontal and vertical polarization components returning from sample and reference paths. Dual silicon photoreceivers measure horizontal and vertical interference fringe intensities versus depth [Γh(z) and Γv(z)].

3.2 Signal conditioning

Detected photocurrents representing Γh(z) and Γv(z) are preamplified, bandpass filtered, and digitized. Coherent demodulation of Γh(z) and Γv(z) yields signals proportional to the horizontal and vertical electric field amplitudes [Eh(z) and Ev(z)] and relative phase [Δϕ(z)] of light backscattered from the specimen at each depth z within the A-scan. An ensemble (NA) of A-scans representing uncorrelated speckle fields are acquired on a grid within a small square region (50 µm×50 µm) at each location of interest on the specimen. Acquisition of an ensemble of NA A-scans at each location is repeated for M incident polarization states distributed uniformly in a great circle on the Poincaré sphere by M phase shifts (δ LCVR,m) of the LCVR. For each M, the calibrated LCVR phase shift (δ LCVR,m) is subtracted from the demodulated relative phase [Δϕm(z)] to compensate for the light’s return propagation through the LCVR. This yields M sets of horizontal and vertical electric field amplitudes [Eh,m(z) and Ev,m(z)] and compensated relative phase [Δϕc,m(z)].

Non-normalized Stokes vectors are calculated from Eh,m(z), Ev,m(z), and Δϕc,m(z) for each of NA A-scans in the ensemble and for each M. Ensemble-averaging over NA at each depth z (denoted by 〈 〉NA) reduces σ speckle by a factor of approximately N 1/2 A and then normalization yields M sets of depth-resolved polarization data [S m(z)] for each location,

Sm(z)=(Q(z)U(z)V(z))=(Eh,m(z)2Ev,m(z)2NA2Eh,m(z)Ev,m(z)cos[Δϕc,m(z)]NA2Eh,m(z)Ev,m(z)sin[Δϕc,m(z)]NA)Eh,m(z)2+Ev,m(z)2NA.
(19)

Wm(z)=(Eh,m(z)2Ev,m(z)22Eh,m(z)Ev,m(z)cos[Δϕc,m(z)]2Eh,m(z)Ev,m(z)sin[Δϕc,m(z)])Eh,m(z)2+Ev,m(z)2NA.
(20)

Wm(z) is used as a scalar weighting factor in the multistate nonlinear algorithm to estimate phase retardation (δ) and relative-attenuation (ε).

3.3 Multistate nonlinear algorithm to determine form-biattenuance

High sensitivity quantification of form-biattenuance (Δχ) is accomplished using a nonlinear fitting algorithm based on the approach introduced by Kemp et al [5

5. N. J. Kemp, J. Park, H. N. Zaatari, H. G. Rylander, and T. E. Milner, “High sensitivity determination of birefringence in turbid media using enhanced polarization-sensitive OCT,” J. Opt. Soc. Am. A 22, 552–560 (2005). [CrossRef]

] for determining form-birefringence (Δn) with PS-OCT. A modified multistate residual function (RM) has been implemented which gives the composite squared deviation between M sets of depth-resolved polarization data [Sm(z)] and corresponding M noise-free model polarization arcs [P m(z)] weighted by Wm(z),

RM=m=1MRo[Sm(zj),Wm(zj);2ε,2δ,β̂,Pm(0)],
(21)

where Ro is the weighted single-state residual function,

Ro=j=1J{W(zj)[S(zj)P[zj;2ε,2δ,β̂,P(0)]}2,
(22)

and the subscript “j” is used to denote the discrete nature of sampled data versus depth (z). Model parameters [2ε, 2δ, β^, and P m(0)] are estimated by minimizing RM using a Levenberg-Marquardt algorithm [19

19. W. H. Press, S. A. Teukolsky, W. T. Vetterling, and B. P. Flannery, Numerical Recipes in C: The Art of Scientific Computing, 2nd Ed. (Cambridge University Press, United Kingdom, 1992).

] and represent the best estimate of P m(z). At increased penetration depths (lower electrical signal-to-noise ratio) or large initial separation-angles [γm(0), as discussed in Section 5.1], Wm(z) decreases and Sm(z) are given less weight. Δχ and Δn are calculated using Eqs. (6) and (4) from estimates of ε and δ provided by the multistate nonlinear algorithm. Δz is measured by subtracting the front and rear specimen boundaries in the OCT intensity image and dividing by the bulk refractive index (n tendon=1.4 and n RNFL=1.38).

Uncertainty in estimates of any single P m(z) arc is offset through constraints placed upon the other M - 1 arcs by the modified residual function [Eq. (21)]. All M noise-free model polarization arcs [P m(z)] must collapse toward the same eigen-axis (β^) at the same rate (2ε) and must rotate around β^ by the same angle (2δ) regardless of the incident polarization state. Discrimination between arc movements on the Poincaré sphere due to either Δn or Δχ is accomplished by restricting contributions from each into orthogonal planes (Sec. 2).

3.4 Ex vivo rat tendon measurements

Four mature, freshly-euthanized Sprague-Dawley rats were obtained from an unrelated study. To collect tail tendon specimens, each tail was cut from the body and a longitudinal incision the length of the tail was made in the skin on the dorsal side. Skin was peeled back and tertiary fascicle groups were extracted with tweezers and placed in phosphate buffered saline solution to prevent dehydration before imaging. Anatomical terminology used in this paper is consistent with that given by Rowe [20

20. R. W. D. Rowe, “The structure of rat tail tendon,” Connect. Tissue Res. 14, 9–20 (1985). [CrossRef] [PubMed]

]. Tertiary fascicle groups were teased apart into individual fascicles with tweezers and placed in a modified cuvette in the sample path of the PS-OCT instrument. The cuvette maintained saline solution around the fascicle, prevented mechanical deformation in the radial direction, and allowed 20 g weights to be attached at each end of the fascicle. Weights provided minimal longitudinal loading in order to flatten the collagen fibril crimp structure present in rat tail tendon [21

21. J. Kastelic, A. Galeski, and E. Baer, “The multicomposite structure of tendon,” Connect. Tissue Res. 6, 11–23 (1978). [CrossRef] [PubMed]

,22

22. S. P. Nicholls, L. J. Gathercole, A. Keller, and J. S. Shah, “Crimping in rat tail tendon collagen: morphology and transverse mechanical anisotropy,” Int. J. Biol. Macromol. 5, 283–88 (1983). [CrossRef]

]. A total of 111 different fascicle locations were imaged from the four rats, each at the location of maximum diameter across its transverse cross-section (as determined by an OCT B-scan image). Achilles tendon specimens from the same rats were harvested in a straightforward manner and imaged in the modified cuvette with the same loading conditions. Four different Achilles tendons were imaged in 45 different randomly chosen locations. Sm(z) was recorded (NA=64, M=3) at all locations and the multistate nonlinear algorithm estimated ε and δ for each location.

To investigate the effect of the relative depth of light focus within the tissue on the estimated values for δ and ε, we imaged the same location on a single rat tail tendon fascicle for a range of axial displacements between the rear principal plane of the f=25 mm focusing lens and the fascicle surface. We recorded Sm(z) (NA=64, M=3) for ten 50-µm-steps from 0 (focused at surface) to 450 µm (focused deep within fascicle).

To investigate the effect of different initial separation-angles [γm(0)] on the PS-OCT-estimated values for δ and ε, we placed a 1/6-wave retarder in the sample path between the LCVR and scanning optics and recorded Sm(z) (NA=64, M=5) from the same location on a single rat tail tendon fascicle for 12 uniformly spaced orientations (between 0° and 165°) of the 1/6-wave retarder axis.

3.5 Ex vivo chicken tendon measurements

We imaged 57 randomly chosen locations in tendons extracted from the proximal end of chicken thighs obtained at a local grocery store. Temperature variations (freezing/thawing or refrigeration) and postmortem time prior to characterization by PS-OCT were unknown. Extracted tendon specimens were kept hydrated in the modified cuvette and imaged without mechanical loading. Sm(z) was recorded (NA=64, M=3) at all locations and the multistate nonlinear algorithm was used to estimate ε and δ for each location.

3.6 In vivo primate retinal nerve fiber layer measurements

Details of the animal protocol for PS-OCT characterization of the in vivo primate retinal nerve fiber layer (RNFL) were given in Ref [18

18. H. G. Rylander, N. J. Kemp, J. Park, H. N. Zaatari, and T. E. Milner, “Birefringence of the primate retinal nerve fiber layer,” Exp. Eye Res. In Press, (2005).

]. Sm(z) was recorded (NA=36, M=6) on two different days for six locations distributed in a 100 µm region around a point 1 mm inferior to the optic nerve head (ONH) center and six locations distributed in a 100 µm region around a point 1 mm nasal to the ONH center.

4. Results

4.1 Rat tail tendon

For 111 locations in rat tail tendon, mean ± standard deviation and [range] in rat tail tendon form-biattenuance were Δχ=5.3·10-4 ± 1.3·10-4 [3.0·10-4, 8.0·10-4] and in form-birefringence were Δn=51.7·10-4 ± 2.6·10-4 [46.8·10-4, 56.3·10-4]. Figures 2(a) and 2(b) show Sm(z) and P m(z) plotted on the Poincaré sphere for two different rat tail tendon fascicles with the largest (Δχ=8.0·10-4) and smallest (Δχ=3.0·10-4) form-biattenuances detected. Form-birefringence for the two fascicles shown in Figures 2(a) and 2(b) were Δn=47.4·10-4 and Δn=55.2·10-4 respectively. Polarimetric signal-to-noise ratio (PSNR) ranged from 51 to 155 and standard deviation of polarimetric speckle noise (σ speckle) was approximately 0.22 rad for the 111 rat tail tendon locations measured.

Fig. 2. Depth-resolved polarization data [S 1(z), orange] and associated noise-free model polarization arc [P 1(z), black] and eigen-axis (β^, green) determined by the multistate nonlinear algorithm in rat tail tendon are shown on the Poincaré sphere (left). Corresponding normalized Stokes parameters [Q(z), U(z), V(z)] and associated nonlinear fits (black) are shown on the right. A single incident polarization state (m=1) is shown for simplicity. (a) Sm(z) for tendon with relatively high form-biattenuance (Δχ=8.0·10-4) collapses toward β^ faster than that for (b) tendon with relatively low form-biattenuance (Δχ=3.0·10-4).

4.2 Variation in form-biattenuance versus relative focal depth

For 10 different displacements between the rear principal plane of the f=25 mm focusing lens and the fascicle surface, σ speckle≈0.22 rad and mean±standard deviation in relative-attenuation were ε=1.42±0.022 rad and in phase retardation were δ=11.5±0.034 rad. Thickness of the tendon specimen was Δz=360 µm.

4.3 Variation in form-biattenuance versus 1/6-wave retarder axis orientation

For 12 orientations of the 1/6-wave retarder axis, mean±standard deviation in relative-attenuation were ε=1.54±0.096 rad and in phase retardation were δ=13.1±0.046 rad. σ speckle increased exponentially from 0.066 rad for a small initial separation-angle of γm(0)=0.81 rad up to σ speckle=0.69 rad for a large γm(0)=3.0 rad. Thickness of the tendon specimen was Δz=383 µm.

4.4 Rat Achilles tendon

For 45 locations in rat Achilles tendon, mean ± standard deviation and [range] in rat Achilles tendon form-biattenuance were Δχ=1.3·10-4±0.53·10-4 [0.74·10-4, 3.2·10-4] and in form-birefringence were Δn=46.9·10-4 ± 5.9·10-4 [32.9·10-4, 56.3·10-4]. Figure 3 shows Sm(z) and P m(z) plotted on the Poincaré sphere for the location in which the form-biattenuance was the lowest of all tendon specimens studied (Δχ=0.74·10-4). PSNR ranged from 76 to 175 and σ speckle≈0.20 rad for the 45 rat Achilles tendon locations measured.

Fig. 3. S1(z) (orange) and associated P 1(z) (black) and β^ (green) determined by the multistate nonlinear algorithm in rat Achilles tendon are shown on the Poincaré sphere (left). A single incident polarization state (m=1) is shown for simplicity. Form-biattenuance in this specimen (Δχ=3.2 °/100µm) is lower than for specimens shown in Figures 2(a) and 2(b) and spiral collapse toward β^ is correspondingly slower.

4.5 Chicken drumstick tendon

For 57 locations in chicken drumstick tendon, mean ± standard deviation and [range] in chicken drumstick tendon form-biattenuance were Δχ=2.1·10-4±0.3·10-4 [1.4·10-4, 3.1·10-4] and in form-birefringence were Δn=44.4·10-4±1.9·10-4 [38.4·10-4, 48.4·10-4]. PSNR ranged from 42 to 96 and σ speckle ≈0.28 rad for the 57 chicken drumstick tendon locations measured.

4.6 In vivo primate RNFL

We report mean ± standard deviation and [range] in form-biattenuance for the six locations within a 100 µm region around a point 1 mm inferior to the ONH center: Δχ=0.18·10-4±0.09·10-4 [0.07·10-4, 0.33·10-4] on day 1 and Δχ=0.18·10-4 ± 0.13·10-4 [0.06·10-4, 0.42·10-4] on day 2. Average RNFL thickness in this region was 166 µm and average relative-attenuation (ε) was 0.023 radians. Figure 4 shows typical Sm(z) and P m(z) plotted on the Poincaré sphere for the region 1 mm inferior to the ONH center in the primate RNFL. PSNR ranged from 3 to 16 and σ speckle ≈ 0.06 rad for the six inferior locations measured. In the region 1 mm nasal to the ONH center, RNFL thickness averaged 50 µm and PSNR was too low for reliable estimates of Δχ in the nasal region of the primate RNFL.

5. Discussion

5.1 Variation in measurements of form-biattenuance

Uncertainty in phase retardation (uδ) was analyzed previously [5

5. N. J. Kemp, J. Park, H. N. Zaatari, H. G. Rylander, and T. E. Milner, “High sensitivity determination of birefringence in turbid media using enhanced polarization-sensitive OCT,” J. Opt. Soc. Am. A 22, 552–560 (2005). [CrossRef]

] and is predominantly due to polarimetric speckle noise (σ speckle) which lingers after ensemble-averaging. Arc length (l arc) has approximately the same functional dependence on δ and ε [Eqs. (17) and (18)]; therefore we expect uncertainty in relative-attenuation (uε) to be similar to uδ for a given σ speckle, though additional experiments in a controlled model are necessary to verify the relationship between uε, uδ, and σ speckle. Uncertainties in form-birefringence (uΔn) or form-biattenuance (χ) are dependent on uδ or uε as well as the specimen thickness (Δz), which complicates comparison of n or χ between specimens or between other variations of PS-OCT. For the rat and chicken tendon specimens we studied (NA=64), σ speckle ranged from 0.20 to 0.28 rad, giving uncertainties (uδ and uε) due to polarimetric speckle noise no higher than ± 0.07 rad. Corresponding uncertainty in form-biattenuance for a Δz=160-µm-thick specimen is χ≈±0.57·10-4. In primate RNFL (NA=36), σ speckle ≈0.06 rad corresponds to uε ≈ ± 0.015 rad or χ≈±0.12·10-4 for an RNFL thickness of Δz=166 µm.

Fig. 4. Sm(z) (colored) and associated P m(z) (black) and β^ (green) for in vivo primate RNFL shown on the Poincaré sphere for M=6 (left). Corresponding normalized Stokes parameters [Q(z), U(z), V(z)] and associated nonlinear fits (black) are shown for a single incident polarization state (m=1, right). Notice the RNFL exhibits only a fraction of a wave of phase retardation compared to multiple waves exhibited by tendon specimens in Figures 2(a), 2(b), and 6. Movie showing 3D nature of Poincaré sphere (1.30 MB).

The range of systematic variation in measurements of δ and ε due to placement of the beam focus was negligible (Sec. 4.2). Variation in measured ε (6.2%) due to different initial separation-angles [γm(0)] was higher than variation in δ (0.35%, Sec. 4.3). Interestingly, σ speckle has a roughly exponential dependence on γm(0). Large initial separation angles [γm(0)≈π] correspond to incident polarization states [Sm(0)] near the preferentially attenuated eigenpolarization; therefore, we expect these Sm(0) to have lower detected intensity and relatively higher noise variation on the Poincaré sphere than Sm(0) with lower γm(0). Additional experiments are needed to characterize completely the dependence of σ speckle on γm(0). Because Wm(z) [Eq. (20)] decreases with increasing σ speckle, states with large γm(0) are weighted less by our multistate nonlinear algorithm when estimating δ and ε. Inspection of the right sides of Figs. 2(a), 2(b), 3, and 4 reveals that σ speckle does not increase significantly versus depth (z) for the limited tissue thicknesses studied. Therefore, we do not expect that reduced collection of light backscattered from deeper in the tissue significantly affects our estimates of ε for the range of depths probed.

5.2 Model for form-biattenuance and form-birefringence

Since the pioneering work of O. Wiener [7

7. O. Wiener, “Die Theorie des Mischkorpers fur das Feld der stationaren Stromung,” Abh. Math.-Phys. Klasse Koniglich Sachsischen Des. Wiss. 32, 509–604 (1912).

], many investigators have examined the origin of form-birefringence in anisotropic media. Notably, Bragg and Pippard [23

23. W. L. Bragg and A. B. Pippard, “The form birefringence of macromolecules,” Acta. Crystallogr. 6, 865–867 (1953). [CrossRef]

] derived an expression for the form-birefringence of a suspension of aligned ellipsoidal particles and compared their result to that derived by Wiener. More recently, Oldenbourg and Ruiz [8

8. R. Oldenbourg and T. Ruiz, “Birefringence of macromolecules: Wiener’s theory revisited, with applications to DNA and tobacco mosaic virus,” Biophys. J. 56, 195–205 (1989). [CrossRef] [PubMed]

] reexamined Wiener’s original theory and applied their model to compute form-birefringence of macromolecules including DNA and the tobacco mosaic virus. None of these investigations address the phenomenon of form-biattenuance nor introduce a model predicting the relative contribution of Δn and Δχ to transformations in polarization state of light propagating in anisotropic media.

In the following, we introduce a rudimentary model for form-biattenuance (Δχ). Consider light normally incident on a fibrous material that consists of alternating anisotropic and isotropic layers (Fig. 5). The anisotropic layer with thickness h 1 is composed of cylindrical fibers (nf) imbedded in water (nw) with center-to-center spacing (a) and with diameters (h 1) much less than the wavelength of incident light (h 1aλ 0). Effective refractive indices parallel (np) and perpendicular (ns) to the fibers are [8

8. R. Oldenbourg and T. Ruiz, “Birefringence of macromolecules: Wiener’s theory revisited, with applications to DNA and tobacco mosaic virus,” Biophys. J. 56, 195–205 (1989). [CrossRef] [PubMed]

]

np2=(h1a)nf2+(1h1a)nw2and
(23)
ns2=nw2+(h1a)(nf2nw2)1+12(1h1a)(nf2nw2nw2).
(24)

The second layer has thickness h - h 1 and contains the same isotropic fluid (water, nw) found between fibers in the adjacent anisotropic layers.

Fig. 5. A model for form-biattenuance consisting of alternating anisotropic and isotropic layers.

When np,s - nw≈0, the Fresnel relations can be used to find expressions for the form-birefringence (Δn) and form-biattenuance (Δχ). The form-birefringence (Δn) is that of the anisotropic layer reduced by the layer fill factor (h 1/h),

Δn=h1h(npns).
(25)

Relative amplitude (tp/ts) between p and s components of normally incident light transmitted to depth z is reduced by transmission through z/h layer-pairs,

tpts(z)=[np(ns+nw)2ns(np+nw)2]zh,
(26)

where multiple reflections between layers are neglected. The form-biattenuance (Δχ) is computed directly from Eqs. (6) and (26) by computing the logarithm,

Δχ=λ02πhln(np(ns+nw)2ns(np+nw)2).
(27)

A notable difference between the functional dependence of Δn and Δχ in this model is apparent in Eqs. (25) and (27). Whereas form-biattenuance (Δχ) depends directly on a wavelength-relative structural dimension (h0), the form of Δn is wavelength independent (ignoring dispersion). Consider for example h 1=0.12 µm collagen fibers (nf=1.51) in water (nw=1.33) with fill factors h 1/a=0.8 and h 1/h=0.8, Eqs. (25) and (27) give Δn=31·10-4 and Δχ=1.2·10-4. Increasing the collagen fiber diameter to h 1=0.18 µm while preserving the fill-factors and material properties gives an identical form-birefringence (Δn=31·10-4) while reducing the form-biattenuance (Δχ=0.81·10-4) by one-third. Interestingly, in this model the ratio of form-biattenuance to form-birefringence (Δχn) is dependent on the structural dimensions of the fibers comprising the material.

Recently Louis-Dorr et al [24

24. V. Louis-Dorr, K. Naoun, P. Alle, A. Benoit, and A. Raspiller, “Linear dichroism of the cornea,” Appl. Opt. 43, 1515–1521 (2004). [CrossRef] [PubMed]

] reported corneal form-biattenuance measurements. That a transparent fibrous biological material such as cornea exhibits form-biattenuance is consistent with the model presented here. Although this model does not include effects such as differential polarimetric scattering and evanescent field propagation which may also lead to form-biattenuance, the computed value (Δχ=1.2·10-4) for 0.12 µm collagen fibers (nf=1.51) in water (nw=1.33) is comparable to that measured experimentally in rat Achilles tendon (Δχ=1.3·10-4). Determining the contribution of differential polarimetric scattering and evanescent field propagation and other candidate mechanisms to the form-biattenuance will require further studies in a model system that allows independent variation of these mechanisms.

5.3 Comparison with previously reported values

The form-biattenuance values we measured in tendon are significantly higher than those approximated from Dz values reported by Park et al [16

16. B. H. Park, M. C. Pierce, B. Cense, and J. F. de Boer, “Jones matrix analysis for a polarization-sensitive optical coherence tomography system using fiber-optic components,” Opt. Lett. 29, 2512–2514 (2004). [CrossRef] [PubMed]

] in chicken tendon (Δχ=0.8·10-4) or by Jiao and Wang [4

4. S. Jiao and L. V. Wang, “Jones-matrix imaging of biological tissues with quadruple-channel optical coherence tomography,” J. Biomed. Opt. 7, 350–358 (2002). [CrossRef] [PubMed]

] in porcine tendon (Δχ=0.17·10-4). The range of Δχ values we measured in a substantial number of specimens of rat tail tendon (3.0·10-4 to 8.0·10-4 for N=111), rat Achilles tendon (0.74·10-4 to 3.2·10-4 for N=45), and chicken drumstick tendon (1.4·10-4 to 3.1·10-4 for N=57) demonstrate that a sizable inter-species and intra-species variation is present in tendon form-biattenuance.

5.4 Relevance and motivation for form-biattenuance

The ability of PS-OCT and the multistate nonlinear algorithm to noninvasively measure form-biattenuance is demonstrated by in vivo imaging of the primate RNFL. We acknowledge that the value we report for form-biattenuance (Δχ=0.18·10-4) in the region 1 mm inferior to the optic nerve head center is marginally higher than our sensitivity (0.12·10-4); however, we can conclude that the reported value represents an upper limit on RNFL form-biattenuance in this region.

Because form-birefringence and form-biattenuance arise from light scattering by nanometer-sized anisotropic structures, development of sophisticated models relating Δn and Δχ to underlying microstructure will allow use of PS-OCT for noninvasively quantifying fibrous constituents (e.g., neurotubules in the RNFL or collagen fibers in tendon) which are smaller than the resolution limit of light microscopy. Although the term form-birefringence is used throughout this paper, PS-OCT is not directly capable of discriminating between form and intrinsic effects; therefore a portion of the reported tendon birefringence may be due to intrinsic birefringence on the molecular scale. Because we expect biattenuance in tendon or RNFL to arise from interactions on the nanometer scale, form-biattenuance and biattenuance are used interchangeably.

Although the small-angle approximation introduces minimal error for previously reported values of diattenuation-per-unit-depth (Dzεz) observed in thin tissue specimens (Δz < 1 mm), we motivate and justify the introduction of a new term, biattenuance (Δχ). First, substantial measurements on tissues studied here have a diattenuation (D) that is outside the range of the small-angle approximation and cannot meaningfully be reported on a diattenuation-per-unit-depth (Dz) basis. Second, biattenuance (Δχ) requires no approximation and is analogous and complementary to a well-understood term, birefringence (Δn). Use of the term biattenuance overcomes the need to clumsily specify when a diattenuation-per-unit-depth approximation is valid or not. Third, consistency in definitions between birefringence (Δn) and biattenuance (Δχ) or between phase retardation (δ) and relative-attenuation (ε) allow a meaningful and intuitive comparison of the relative values (i.e. Δχn, ε/δ) of amplitude and phase anisotropy in any optical medium or specimen. Fourth, availability of narrow line-width swept-source lasers may allow construction of Fourier-domain PS-OCT instruments having scan depths far longer than current PS-OCT instruments. By using these sources and hyperosmotic agents to reduce scattering in tissue [25

25. G. Vargas, E. K. Chan, J. K. Barton, H. G. Rylander, and A. J. Welch, “Use of an agent to reduce scattering in skin,” Lasers in Surgery and Medicine 24, 133–141 (1999). [CrossRef] [PubMed]

], future studies using PS-OCT may probe significantly deeper into tissue specimens than 1–2 mm, likely making the small-angle approximation invalid even in tissues with low biattenuance. Fifth, PS-OCT may be applied to characterize non-biological samples [26

26. K. Wiesauer, M. Pircher, E. Goetzinger, S. Bauer, R. Engelke, G. Ahrens, G. Grutzner, C. K. Hitzenberger, and D. Stifter, “En-face scanning optical coherence tomography with ultra-high resolution for material investigation,” Opt. Express 13, 1015–1024 (2005), http://www.opticsexpress.org/abstract.cfm?URI=OPEX-13-3-1015. [CrossRef] [PubMed]

] which may have high D and not satisfy the small-angle approximation. Sixth, biattenuance may be useful to investigators employing other polarimetric optical characterization techniques which can detect anisotropically scattered light and for which dichroism is therefore inappropriate. Finally, although the term “depth-resolved” is frequently used in OCT literature in the context of either “measured in the depth dimension” or “local variation in a parameter versus depth [e.g., Δχ(z)]”, motivation for biattenuance is independent of the particular interpretation. The first interpretation is assumed for work reported here, but the multistate nonlinear algorithm can be extended in a straightforward manner to provide local variation in biattenuance versus depth [Δχ(z)].

6. Conclusion

A summary of the terms and symbols used in this paper is appropriate here. Biattenuance (Δχ) is an intrinsic physical property responsible for polarization-dependent amplitude attenuation, just as birefringence (Δn) is the physical property responsible for polarization-dependent phase delay. Diattenuation (D) gives the quantity of accumulated anisotropic attenuation over a given depth (Δz) by a given optical element. The nonlinear dependence of diattenuation on depth motivated our introduction of relative-attenuation (ε), which depends linearly on depth, maintains parallelism and consistency with phase retardation (δ) in Eq. (7), and is a natural parameter in depth-resolved polarimetry such as PS-OCT. The mathematical relationships between these parameters were given in Eqs. (4), (6), and (8).

Acknowledgments

The authors gratefully acknowledge the Optics Express editor and reviewers for helpful suggestions to improve this paper, and the Texas Higher Education Coordinating Board Advanced Technology Program (003658-0359-1999), the National Eye Institute at NIH (R24EY12877-02), and the NSF-IGERT program for financial support of this research. Thanks also to Jennifer Cassaday for procurement of rat specimens.

References and links

1.

C. K. Hitzenberger, E. Gotzinger, M. Sticker, M. Pircher, and A. F. Fercher, “Measurement and imaging of birefringence and optic axis orientation by phase resolved polarization sensitive optical coherence tomography,” Opt. Express 9, 780–790 (2001), http://www.opticsexpress.org/abstract.cfm?URI=OPEX-9-13-780. [CrossRef] [PubMed]

2.

B. H. Park, C. Saxer, T. Chen, S. M. Srinivas, J. S. Nelson, and J. F. de Boer, “In vivo burn depth determination by high-speed fiber-based polarization sensitive optical coherence tomography,” J. Biomed. Opt. 6, 474–479 (2001). [CrossRef] [PubMed]

3.

M. G. Ducros, J. D. Marsack, H. G. Rylander, S. L. Thomsen, and T. E. Milner, “Primate retinal imaging with polarization-sensitive optical coherence tomography,” J. Opt. Soc. Am. A 18, 2945–2956 (2001). [CrossRef]

4.

S. Jiao and L. V. Wang, “Jones-matrix imaging of biological tissues with quadruple-channel optical coherence tomography,” J. Biomed. Opt. 7, 350–358 (2002). [CrossRef] [PubMed]

5.

N. J. Kemp, J. Park, H. N. Zaatari, H. G. Rylander, and T. E. Milner, “High sensitivity determination of birefringence in turbid media using enhanced polarization-sensitive OCT,” J. Opt. Soc. Am. A 22, 552–560 (2005). [CrossRef]

6.

M. Born and E. Wolf, Principles of Optics (Cambridge University Press, Cambridge, 1959).

7.

O. Wiener, “Die Theorie des Mischkorpers fur das Feld der stationaren Stromung,” Abh. Math.-Phys. Klasse Koniglich Sachsischen Des. Wiss. 32, 509–604 (1912).

8.

R. Oldenbourg and T. Ruiz, “Birefringence of macromolecules: Wiener’s theory revisited, with applications to DNA and tobacco mosaic virus,” Biophys. J. 56, 195–205 (1989). [CrossRef] [PubMed]

9.

R. A. Chipman, “Polarization analysis of optical systems,” Opt. Eng. 28, 90–99 (1989).

10.

S.-Y. Lu and R. A. Chipman, “Interpretation of Mueller matrices based on polar decomposition,” J. Opt. Soc. Am. A 13, 1106–1113 (1996). [CrossRef]

11.

R. M. Craig, S. L. Gilbert, and P. D. Hale, “Accurate Polarization Dependent Loss Measurement and Calibration Standard Development,” Symposium on Optical Fiber Measurements NIST Special Publication 930, 5–8 (1998).

12.

B. Huttner, C. Geiser, and N. Gisin, “Polarization-Induced Distortions in Optical Fiber Networks with Polarization-Mode Dispersion and Polarization-Dependent Losses,” IEEE J. Sel. Top. Quantum Electron. 6, 317–329 (2000). [CrossRef]

13.

J. W. Verhoeven, “Glossary of terms used in photochemistry,” Pure App. Chem. 68, 2228 (1996).

14.

J. F. de Boer, T. E. Milner, and J. S. Nelson, “Determination of the depth-resolved Stokes parameters of light backscattered from turbid media using Polarization Sensitive Optical Coherence Tomography,” Opt. Lett. 24, 300–302 (1999). [CrossRef]

15.

M. Todorovic, S. Jiao, L. V. Wang, and G. Stoica, “Determination of local polarization properties of biological samples in the presence of diattenuation by use of Mueller optical coherence tomography,” Opt. Lett. 29, 2402–2404 (2004). [CrossRef] [PubMed]

16.

B. H. Park, M. C. Pierce, B. Cense, and J. F. de Boer, “Jones matrix analysis for a polarization-sensitive optical coherence tomography system using fiber-optic components,” Opt. Lett. 29, 2512–2514 (2004). [CrossRef] [PubMed]

17.

J. Park, N. J. Kemp, H. N. Zaatari, H. G. Rylander III, and T. E. Milner, Dept. of Biomedical Engineering, University of Texas, 1 University Station, #C0800, Austin, TX 78712 are preparing a manuscript to be called “Differential geometry of the normalized Stokes vector trajectories in anisotropic media.”

18.

H. G. Rylander, N. J. Kemp, J. Park, H. N. Zaatari, and T. E. Milner, “Birefringence of the primate retinal nerve fiber layer,” Exp. Eye Res. In Press, (2005).

19.

W. H. Press, S. A. Teukolsky, W. T. Vetterling, and B. P. Flannery, Numerical Recipes in C: The Art of Scientific Computing, 2nd Ed. (Cambridge University Press, United Kingdom, 1992).

20.

R. W. D. Rowe, “The structure of rat tail tendon,” Connect. Tissue Res. 14, 9–20 (1985). [CrossRef] [PubMed]

21.

J. Kastelic, A. Galeski, and E. Baer, “The multicomposite structure of tendon,” Connect. Tissue Res. 6, 11–23 (1978). [CrossRef] [PubMed]

22.

S. P. Nicholls, L. J. Gathercole, A. Keller, and J. S. Shah, “Crimping in rat tail tendon collagen: morphology and transverse mechanical anisotropy,” Int. J. Biol. Macromol. 5, 283–88 (1983). [CrossRef]

23.

W. L. Bragg and A. B. Pippard, “The form birefringence of macromolecules,” Acta. Crystallogr. 6, 865–867 (1953). [CrossRef]

24.

V. Louis-Dorr, K. Naoun, P. Alle, A. Benoit, and A. Raspiller, “Linear dichroism of the cornea,” Appl. Opt. 43, 1515–1521 (2004). [CrossRef] [PubMed]

25.

G. Vargas, E. K. Chan, J. K. Barton, H. G. Rylander, and A. J. Welch, “Use of an agent to reduce scattering in skin,” Lasers in Surgery and Medicine 24, 133–141 (1999). [CrossRef] [PubMed]

26.

K. Wiesauer, M. Pircher, E. Goetzinger, S. Bauer, R. Engelke, G. Ahrens, G. Grutzner, C. K. Hitzenberger, and D. Stifter, “En-face scanning optical coherence tomography with ultra-high resolution for material investigation,” Opt. Express 13, 1015–1024 (2005), http://www.opticsexpress.org/abstract.cfm?URI=OPEX-13-3-1015. [CrossRef] [PubMed]

OCIS Codes
(030.6140) Coherence and statistical optics : Speckle
(170.3880) Medical optics and biotechnology : Medical and biological imaging
(170.4500) Medical optics and biotechnology : Optical coherence tomography
(260.1440) Physical optics : Birefringence
(260.5430) Physical optics : Polarization

ToC Category:
Research Papers

History
Original Manuscript: March 22, 2005
Revised Manuscript: April 22, 2005
Published: June 13, 2005

Citation
Nate Kemp, Haitham Zaatari, Jesung Park, H. Grady Rylander III, and Thomas Milner, "Form-biattenuance in fibrous tissues measured with polarization-sensitive optical coherence tomography (PS-OCT)," Opt. Express 13, 4611-4628 (2005)
http://www.opticsinfobase.org/oe/abstract.cfm?URI=oe-13-12-4611


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References

  1. C. K. Hitzenberger, E. Gotzinger, M. Sticker, M. Pircher and A. F. Fercher, "Measurement and imaging of birefringence and optic axis orientation by phase resolved polarization sensitive optical coherence tomography," Opt. Express 9, 780-790 (2001), <a href="http://www.opticsexpress.org/abstract.cfm?URI=OPEX-9-13-780.">http://www.opticsexpress.org/abstract.cfm?URI=OPEX-9-13-780.</a> [CrossRef] [PubMed]
  2. B. H. Park, C. Saxer, T. Chen, S. M. Srinivas, J. S. Nelson and J. F. de Boer, "In vivo burn depth determination by high-speed fiber-based polarization sensitive optical coherence tomography," J. Biomed. Opt. 6, 474-479 (2001). [CrossRef] [PubMed]
  3. M. G. Ducros, J. D. Marsack, H. G. Rylander, S. L. Thomsen and T. E. Milner, "Primate retinal imaging with polarization-sensitive optical coherence tomography," J. Opt. Soc. Am. A 18, 2945-2956 (2001). [CrossRef]
  4. S. Jiao and L. V. Wang, "Jones-matrix imaging of biological tissues with quadruple-channel optical coherence tomography," J. Biomed. Opt. 7, 350-358 (2002). [CrossRef] [PubMed]
  5. N. J. Kemp, J. Park, H. N. Zaatari, H. G. Rylander and T. E. Milner, "High sensitivity determination of birefringence in turbid media using enhanced polarization-sensitive OCT," J. Opt. Soc. Am. A 22, 552-560 (2005). [CrossRef]
  6. M. Born and E. Wolf, Principles of Optics (Cambridge University Press, Cambridge, 1959).
  7. O. Wiener, "Die Theorie des Mischkorpers fur das Feld der stationaren Stromung," Abh. Math.-Phys. Klasse Koniglich Sachsischen Des. Wiss. 32, 509-604 (1912).
  8. R. Oldenbourg and T. Ruiz, "Birefringence of macromolecules: Wiener's theory revisited, with applications to DNA and tobacco mosaic virus," Biophys. J. 56, 195-205 (1989). [CrossRef] [PubMed]
  9. R. A. Chipman, "Polarization analysis of optical systems," Opt. Eng. 28, 90-99 (1989).
  10. S.-Y. Lu and R. A. Chipman, "Interpretation of Mueller matrices based on polar decomposition," J. Opt. Soc. Am. A 13, 1106-1113 (1996). [CrossRef]
  11. R. M. Craig, S. L. Gilbert and P. D. Hale, "Accurate Polarization Dependent Loss Measurement and Calibration Standard Development," Symposium on Optical Fiber Measurements NIST Special Publication 930, 5-8 (1998).
  12. B. Huttner, C. Geiser and N. Gisin, "Polarization-Induced Distortions in Optical Fiber Networks with Polarization-Mode Dispersion and Polarization-Dependent Losses," IEEE J. Sel. Top. Quantum Electron. 6, 317-329 (2000). [CrossRef]
  13. J. W. Verhoeven, "Glossary of terms used in photochemistry," Pure App. Chem. 68, 2228 (1996).
  14. J. F. de Boer, T. E. Milner and J. S. Nelson, "Determination of the depth-resolved Stokes parameters of light backscattered from turbid media using Polarization Sensitive Optical Coherence Tomography," Opt. Lett. 24, 300-302 (1999). [CrossRef]
  15. M. Todorovic, S. Jiao, L. V. Wang and G. Stoica, "Determination of local polarization properties of biological samples in the presence of diattenuation by use of Mueller optical coherence tomography," Opt. Lett. 29, 2402-2404 (2004). [CrossRef] [PubMed]
  16. B. H. Park, M. C. Pierce, B. Cense and J. F. de Boer, "Jones matrix analysis for a polarization-sensitive optical coherence tomography system using fiber-optic components," Opt. Lett. 29, 2512-2514 (2004). [CrossRef] [PubMed]
  17. J. Park, N. J. Kemp, H. N. Zaatari, H. G. Rylander III, and T. E. Milner, Dept. of Biomedical Engineering, University of Texas, 1 University Station, #C0800, Austin, TX 78712 are preparing a manuscript to be called �??Differential geometry of the normalized Stokes vector trajectories in anisotropic media.�??
  18. H. G. Rylander, N. J. Kemp, J. Park, H. N. Zaatari and T. E. Milner, "Birefringence of the primate retinal nerve fiber layer," Exp. Eye Res. In Press, (2005).
  19. W. H. Press, S. A. Teukolsky, W. T. Vetterling and B. P. Flannery, Numerical Recipes in C: The Art of Scientific Computing, 2nd Ed. (Cambridge University Press, United Kingdom, 1992).
  20. R. W. D. Rowe, "The structure of rat tail tendon," Connect. Tissue Res. 14, 9-20 (1985). [CrossRef] [PubMed]
  21. J. Kastelic, A. Galeski and E. Baer, "The multicomposite structure of tendon," Connect. Tissue Res. 6, 11-23 (1978). [CrossRef] [PubMed]
  22. S. P. Nicholls, L. J. Gathercole, A. Keller and J. S. Shah, "Crimping in rat tail tendon collagen: morphology and transverse mechanical anisotropy," Int. J. Biol. Macromol. 5, 283-88 (1983). [CrossRef]
  23. W. L. Bragg and A. B. Pippard, "The form birefringence of macromolecules," Acta. Crystallogr. 6, 865-867 (1953). [CrossRef]
  24. V. Louis-Dorr, K. Naoun, P. Alle, A. Benoit and A. Raspiller, "Linear dichroism of the cornea," Appl. Opt. 43, 1515-1521 (2004). [CrossRef] [PubMed]
  25. G. Vargas, E. K. Chan, J. K. Barton, H. G. Rylander and A. J. Welch, "Use of an agent to reduce scattering in skin," Lasers in Surgery and Medicine 24, 133-141 (1999). [CrossRef] [PubMed]
  26. K. Wiesauer, M. Pircher, E. Goetzinger, S. Bauer, R. Engelke, G. Ahrens, G. Grutzner, C. K. Hitzenberger and D. Stifter, "En-face scanning optical coherence tomography with ultra-high resolution for material investigation," Opt. Express 13, 1015-1024 (2005), <a href="http://www.opticsexpress.org/abstract.cfm?URI=OPEX-13-3-1015.">http://www.opticsexpress.org/abstract.cfm?URI=OPEX-13-3-1015.</a> [CrossRef] [PubMed]

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