Imaging of arterial plaque by quadrature swept-source optical coherence tomography with signal to noise ratio enhancements

doi:10.4236/jbise.2011.412093

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J. Biomedical Science and Engineering, 2011, 4, 755-761

doi:10.4236/jbise.2011.412093 Published Online December 2011 (http://www.SciRP.org/journal/jbise/

JBiSE

Published Online December 2011 in SciRes. http://www.scirp.org/journal/JBiSE

Imaging of arterial plaque by quadrature swept-source optical

coherence tomography with signal to noise ratio enhancements

Youxin Mao*, Costel Flueraru, Shoude Chang

Institute for Microstructural Sciences, National Research Council Canada, Ottawa, Canada.

Email: *[email protected]

Received 18 August 2011; revised 5 October 2011; accepted 11 November 2011.

ABSTRACT

Arterial plaque from a myocardial infarction-prone

Watanabe heritable hyperlipidemic (WHHLMI) rab-

bit is visualized and characterized using a signal to

noise ratio enhanced swept-source optical coherence

tomography system with a quadrature interferometer

(QSS-OCT). A semiconductor optical amplifier is

used in the sample arm to amplify the weak signal

scattered from arterial plague. Signal to noise ratio

improvement are demonstrated in our QSS-OCT

system. This finding results into an increase of the

penetration depth possible in OCT images, from 1

mm to 2 mm. Preliminary results show that vulner-

able plaque with fibrous cap, macrophage accumula-

tions and calcification in the arterial tissue are meas-

urable with our QSS-OCT system.

Keywords: Optical Coherence Tomography; Arterial

Plaque; Medical and Biological Imaging; Semiconductor

Optical Amplifier

1. INTRODUCTION

The identification of unstable plaque is central in risk-

stratifying patients for acute coronary events. Optical

coherence tomography (OCT) [1] is a recently intro-

duced imaging modality that has shown considerable

promise for the identification of high-risk plaques. The

advantages of OCT compared to ultrasound include its

higher resolution, video speed acquisition rate, com-

pactness and portability. When a small and inexpensive

optical fiber probe as an optical catheter constitutes the

sample arm, the system becomes suitable for intra-vas-

cular probing [2]. Because OCT uses light, a variety of

functional and spectroscopic techniques are available to

expand its capabilities, including polarization, absorp-

tion, elastography, Doppler, and dispersion analysis.

An OCT system with higher signal-to-noise ratio

(SNR) is essentially important for imaging turbid tissues,

such as arterial plaques, because the backscattered sig-

nals from these types of samples are extremely weak.

Swept-source OCT (SS-OCT) has received much atten-

tion in recent years not only because of its higher SNR at

high imaging speeds but also for its imaging possibilities

in the 1300 nm wavelength range, where the reduced

light scattering by tissue enables OCT to collect signal

from deeper into tissue compared to OCT imaging based

on shorter source wavelengths. SS-OCT could also use

the quadrature interferometer based on multi-port fiber

couplers such as the 3 × 3 quadrature interferometer [3].

When measuring instantaneous complex signals with

stable phase information by using a 3 × 3 quadrature

interferometer one can suppress the complex conjugate

artefact and therefore double the effective imaging depth

[4,5]. The phase information of the complex interfer-

ometric signals can also be exploited to gain additonal

information about the tissue, to enhance image contrast

and to perform quantitative measurements. In addition,

the Mach-Zehnder configuration of the presented inter-

ferometric setup (MZI) allows different options to dis-

tribute optical power between the reference and sample

arms. An unbalanced input directs more optical power

from the light source to the sample than that to the ref-

erence mirror [6] while the balanced detection is used to

reduce the beat noise [7]. Both techniques play their

parts in increasing SNR. However in the case of imaging

turbid biomedical samples, the signal backscattered from

tissue is much weaker than the reference signal so an

attenuation of optical power in the reference arm is re-

quired in order to increase the SNR. The ability of OCT

to image vascular plaque has been previously demon-

strated [8-11]. However, OCT images of the arterial wall

are limited to depths of ~1 mm even using the Fourier

domain methods [2]. The limited imaging depth into the

vascular wall is one of the most serious limitations for

OCT to be used as a routine clinical intravascular imag-

ing method. Further improvement of the SNR of OCT is

needed in order to increase the imaging depth of OCT

into tissue. Adding an optical amplifier in the path of the

backscattered signal in the sample arm of an SS-OCT

Y. X. Mao et al. / J. Biomedical Science and Engineering 4 (2011) 755-761

756

system with a balanced Michelson interferometer [12]

and a 2 × 2 MZI configuration [13,14] has been pro-

posed. Modest amount improvement of signal to noise

ratio or sensitivity had reported in their configurations.

However, in our knowledge, no OCT image improve-

ment in cardiology applications has been reported.

In this paper, we theoretically and experimentally

demonstrate a SS-OCT system with a quadrature inter-

ferometer (QSS-OCT) using a semiconductor optical

amplifier (SOA) for the amplification of the weak signal

existent in the sample arm. Improvement of SNR is

demonstrated. Lipid-rich plaque of a WHHLMI rabbit is

visualized and characterized with this system. Prelimi-

nary results show that vulnerable plaque with fibrous cap,

macrophage accumulations and calcification in the arte-

rial tissue are measurable with our QSS-OCT system,

which is also able to image features located as deep as 2

mm from the lumen surface.

2. METHODS AND MATERIALS

Figure 1 shows an experimental setup of our QSS-OCT

system that uses a balanced 3 × 3 and 2 × 2 quadrature

MZI and a SOA for weak sample signal amplification.

The swept laser source (HSL2000-HL, Santac) used in

the setup had a central wavelength of 1320 nm and a full

scan wavelength range of 110 nm, was swept linearly in

optical frequency with a linearity of 0.2%. The band-

width of the source corresponds to an 8-μm OCT imag-

ing resolution in the air. The average output power and

coherence length of the swept source was 12 mW and 10

mm, respectively. A repetition scan rate of 20 kHz was

used in our system and the related duty cycle was 68%.

The light output from the swept laser source was launched

first into the 2 × 2 coupler where 90% of the power was

diverted toward the sample. The reference arm was ar-

ranged with a fiber collimator and a mirror. The light

was directed to the sample through a lensed single mode

fiber probe [15]. A galvanometer (Blue Hill Optical

technologies) scanned the fiber probe along the sample

surface up to 4 mm-long trip corresponding to an OCT

image width of 900 pixels. The weak light back- scat-

tered from the sample was fed into a SOA (Covega)

whose gain can be adjusted by a variable attenuator

connected after the SOA. The SOA had the same center

wavelength and bandwidth as the swept source. The gain

could be varied from 15 dB to 35 dB. The reference arm

has been build so that it can match the optical distance of

the sample arm without SOA and by adding an optical

jumper with certain length it can match the optical dis-

tance of the sample arm with SOA. Both, the SOA and

the added optical jumper (part of reference arm) could

be removed allowing the system to be switched back to a

regular QSS-OCT system without sample signal ampli-

fication. A polarization controller was inserted before

the SOA for optimal amplification. The amplified signal

was combined with the signal returning from the refer-

ence mirror through the 3 × 3 and 2 × 2 couplers, thus

implementing a dual-channel balanced detection system

with two complementary components of the complex

interferometric signals which suppresses the complex

conjugate artefact. Both balanced detectors (PDB150C,

Thorlabs) used in this system had saturation powers of 5

mW. We selected a 3 dB bandwidth of 50 MHz to give

sufficient imaging depth. The two detector outputs were

Figure 1. Experimental setup of our QSS-OCT system with a balanced 3 × 3 and 2 × 2 quad-

rature MZI and an SOA for the amplification of the signal back-scattered from the sample.

Y. X. Mao et al. / J. Biomedical Science and Engineering 4 (2011) 755-761 757

digitized using a data acquisition card (DAQ) (Ala-

zartech, Montreal) with 14-bit resolution and acquired

signal at a sampling speed of 100 MS/s. The swept

source generated a start trigger signal that was used to

initiate the function generator for the galvo scanner and

initiate the data acquisition process for each A-scan.

Because the swept source was linearly swept with the

wave-number k, A-scans data with resolved complex

conjugate artifact were obtained by a direct inverse Fou-

rier transformation (IFT) from the DAQ sampled data

without performing an additional re-sampling step.

Watanabe heritable hyperlipidemic (WHHLMI) rabbit

is a suitable animal model to study familial hypercho-

lesterolemia and atherosclerosis. Arteries in these rabbits

develop atheromatous plaques similar to those in hu-

mans [16]. Figure 2(a) shows a picture of a segment of

the descending aorta together with the protected for-

ward-view ball fiber probe collecting an OCT image in

this work. The size of the OCT probe and the properties

of the probing beam are important for OCT imaging. An

optical fiber-lensed probe with a diameter as small as 0.5

mm is suitable for OCT imaging, especially for in vivo

intravascular imaging. Based on interaction of near in-

frared light with different human tissues, the range of

penetration depth is from 0.5 mm to 3 mm [17]. Work-

ing distance range of an ultra-small fiber-based lense [15]

can be designed from 0.4 to 1.2 mm for matching the

penetration depth of tissues tested. Depth of field can be

in the range of 0.3 - 1.5 mm, which corresponds the spot

size range of 15 - 35 μm at the 1300 nm wavelength,

because the tradeoff between the depth of field and beam

spot size for a Gaussian beam. For imaging arterial tis-

sue, a fiber ball lens was designed and fabricated in

house with a ball size of 0.3 mm, the working distance

of 1.25 mm, depth of field of 1.0 mm, and 1/e2 spot di-

ameter of 29 μm shown in the inset of Figure 2(a) as a

forward—view probe. To form a side-view fiber probe,

the output beam can be total internal reflected 90 - 100

degree by a 45 - 50 degree polished face on the fiber ball.

As a sample, a needle delivered fiber catheter probe de-

signed for the OCT intravascular imaging is shown in

Figure 2(b). The polished lens and the uncoated portion

of the SMF are protected in a transparent inner catheter

(OD 0.49 mm) shown in the inset of Figure 2(b). The

buffered portion of the fiber is attached to an outer flexi-

ble catheter (OD 1.4 mm) after a syringe, which is fas-

tened onto a modified syringe piston (not shown here),

while the transparent inner catheter is inserted into a 21

G (OD 0.81 mm) echogenic spinal needle (VWR, Mis-

sissauga, ON, Canada). After insertion into the tissue,

(a)

(b)

Figure 2. (a) An opened left descending coronary tissue from a

Watanabe heritable hyperlipidemic (WHHLMI) rabbit being

scanned with the forward ball lens fiber probe protected by a

plastic tube. Inset: A scanning electron micrograph of the fiber

ball lens with forward-view fabricated in our lab. (b) A needle

tip of a side-view fibber ball lens probe as a sample of an OCT

probe for in vivo intravascular imaging.

the needle can be drawn back a small distance to let the

optical probe expose to the tissue as shown in Figure

2(b). The probe is then scanned axially inside the tissue

driven by a linear scanner, such that a two dimensional

OCT image is formed.

3. SIGNAL TO NOISE RATIO ANALYSIS

To estimate SNR in our QSS-OCT system shown in

Figure 1(a), assuming that the signal in the sample arm

is coming from a single layer reflector located at the

depth of z0, the two channel currents on the positive and

negative photodiodes in each balanced detector when the

SOA is inserted into the sample arm are given by [14,

18]:

 



 

1 0

,,2,cos

mrmsmSPrmsm m

IkGPkPkGP GPkPkGkz













 



 (1)

Y. X. Mao et al. / J. Biomedical Science and Engineering 4 (2011) 755-761

758

 



 

2 0

,,2,cos2

mrmsmSPrmsm m

IkGPkPkGP GPkPkGkz











 







(2)

where,



ehv



 is photodiode conversion factor,



is the quantum efficiency of the detector, e is the elec-

tronic charge, h is the Planck’s constant,



0 is the mean

frequency of the incident light, is the optical

signal power from the reference arm arriving at the pho-

todiodes when wavenumber k = km, is the

amplified optical signal power from the sample arm at

the photodiodes when wavenumber k = km,





Pk G

complex discrete Fourier transform (DFT). Then, the ma-

ximum-squared signal power in our system with the dual-

balanced quadrature detection is given as [18]:

  

4rs



GDPP



G (3)







P is

the spontaneous emission power of the SOA recorded at

the photodiodes, G is the gain of SOA,



is an arbi-

trary phase shift,



 is the phase shift between the two

output signals, m = 1, 2, , M, where M is the total

sampling number of the axial pixels. Balanced detection

subtracts each positive and negative current for each

channel shown in Eqs.1 and 2, so that the common DC

parts are subtracted and the opposite AC parts are multi-

plied. The complementary phase components of the sig-

nals can be calculated by Eqs.8 and 9 i

n Ref. [4] for a



where the brackets denote the ensemble average, Pr is

the optical power impinging on each photodiode re-

flected from the reference mirror.





PG is the ampli-

fied signal power incident on each photodiode backscat-

tered from the sample. D is a multiplied factor coming

from complex IFT, with D = 2 in our setup.

In this configuration there are three typical noise

sources: thermal noise, shot noise and beat noise [19].

By extending the noise analysis of SS-OCT [18] with

slightly mismatched balanced detection [20] to our QSS-

OCT system with the SOA inserted into on the sample

arm, the total noise power as a function of the SOA gain,

G, is obtained:

 



 





 

222

22 2

4222 1

thrSPb rsprrs sp

GDBiePPGRINPP GRINPRINP G

 





 







 (4)

where, B is the detector bandwidth, ith is the thermal

current of the detector, β is a balanced factor [21] with β =

1 representing a balanced system, RINr/s = 2/δνr/s is the

relative intensity noise of source and SOA, δνr and δνs is

their effective bandwidth, respectively. RINb is the rela-

tive cross-beat intensity noise from the reference light

and the spontaneous emission power of the SOA, which

can be estimated from the experimental results. The first

and second terms in Eq.4 are thermal noise and shot noise

as commonly expressed. The third term represents the

beat noise, which includes the cross-beat and self-beat

noise of the two arms. If the losses and reflectivities of

the reference and sample arms are defined as γr, Rr and γs,

Rs, then the powers from the reference and sample arms

impinging on the photodiodes can be described as PoγrARr

and PoγsGRs, respectively, where A is the reference at-

tenuation when the sample signal is not being amplified.

If the loss from the SOA to the detector is defined as γ,

the spontaneous emission power of the SOA impinging

on the photodiodes can be written as γPsp. The SNR of

our QSS-OCT system with the SOA inserted for weak

sample signal amplification is described as:

 







 





22 2

det 0

22 22

2221

rrss

rr sp

brrsprrrs sp

DP ARGR

SNR G

BiePARP G

RINP ARPGRINP ARRINPG

 

 









 

 

 

 



(5)

Figure 3 shows the calculated results for various

noise powers (a) and SNR (b) versus SOA gain, G, when

the QSS-OCT system contains a sample signal amplifier

(right-side horizontal axis) and versus reference arm

attenuation, A, when the QSS-OCT system without a

sample signal amplifier (left-side horizontal axis). The

calculations are based on the following assumptions: λ0

= 1.32 μm, δλr = 1 nm, P0 = 12 mW, γr = –20 dB, γs =

–10.5 dB, γ = –9 dB, Rr = 0 dB, Rs = –55 dB, B = 50

MHz, β = 0.99, ith = 2 n A/sqrt (Hz), δλs = 100 nm, Psp

= Psp0G, Psp0 = 10–3 mW.

From the theoretical analysis results shown in Figure

3, when the QSS-OCT system without the SOA, refer-

ence light power could be attenuated to reduce the ref-

erence power self-beat noise and to obtain the shot-noise

limit. When the QSS-OCT system contains the SOA on

the sample arm, the SNR can be increased as the gain of

SOA increase although the system is no longer shotnoise

Y. X. Mao et al. / J. Biomedical Science and Engineering 4 (2011) 755-761 759

(a)

(b)

Figure 3. Theoretical analysis results of various noise powers

(a) and SNR (b) versus SOA gain, G, when the QSS-OCT

system contains a sample signal amplifier (right-side horizontal

axis) and versus reference arm attenuation, A, when the QSS-

OCT system without a sample signal amplifier (left-side hori-

zontal axis).

limit. When the gain of sample arm SOA is low, < 15 dB,

reference (ref.) self-beat noise dominates the cross-beat

noise, the SNR linearly increases as the gain is increased

because the ref. self-beat noise stays constant as the gain

increases. Because the cross-beat noise from the refer-

ence power and the spontaneous power emission of the

SOA increases as the gain increases, when G > 15 dB,

the cross-beat noise raises to the level of the ref.

self-beat noise, increase of SNR becomes slowly. As the

gain continuously increases to G > 30 dB, SNR will

saturate where the cross-beat noise becomes dominant.

Increase of SNR up to 18 dB is calculated when the

QSS-OCT system contains the optical amplifier in com-

parison with the system without the optical amplifier in

the shot noise limit as shown in Figure 3.

4. RESULTS AND DISCUSSION

Figure 4(a) shows an ex vivo OCT image of a segment

of aorta where the image was acquired using the

QSS-OCT system with a 25 dB gain SOA inserted in

sample arm. The axial and lateral dimensions of the im-

age are 500 × 900 pixels respectively which correspond

to an image size of 2 × 3 mm2. The image size is cali-

brated using a 1 mm glass slice and assume refractive

index of arterial is 1.3. In Figure 4(a), a clear raised lipid

core with macrophage accumulation with un-uniform

reflectance within its volume (black arrows) is shown. A

thin fibrous cap, which strongly scatters light, covers the

lipid core (white circle). The fibrous cap was defined as

the minimum distance from the coronary artery lumen to

the upper border of the lipid pool. A few calcified re-

gions (white arrows) around the lipid core, characterized

itself by low reflectance, are clearly discernable. The

regions with uniform reflectance (grey arrows) corre-

spond to bundles of smooth muscle cells, which can be

viewed up to depths of 2 mm even beneath the calcified

regions. The results obtained from images of WHHLMI

coronaries acquired with our QSS-OCT, as the image

shown in Figure 4(a), agreed very well with the histo-

logical micrographs of these samples, shown in the Fig-

ure 4 of ref. [16]. For comparison, an OCT image ac-

quired at the same position on the sample after the SOA

was extracted from the sample arm is shown in Figure

4(b). Obviously, the image shown in Figure 4(b) does

not provide a clear view of all the clinical aspects of the

sample, especially in the regions located deeper than 1

mm. To further quantify the analysis, we selected three

A-scan profiles which are illustrated in Figures 4(c)-(e).

These intensity profiles are shown as they were acquired

with the QSS-OCT system when it had an SOA (black)

inserted and without SOA inserted (grey). These three

scans are located at the positions marked with dashed

arrows in Figures 4(a) and (b). By comparing the data

recorded when the system had the SOA with the data

recorded by the system without SOA, it can be observed

that the signal values increase up to 25 dB while the lev-

els of noise increase by only 10 dB, so that SNR of 15

dB is increased. Signals coming from structures located

at depths of up to 2 mm can be observed in the data ac-

quired with the system that has the SOA inserted. How-

ever, structures located at and deeper than 1 mm become

difficult to distinguish when the SOA is removed. The

increase in the image penetration depth when acquired

by the QSS-OCT with the SOA inserted is also evident

in Figure 4(a) from the ability to distinguish the calcifi-

cation boundary. The identified features from Figure 4(a)

can be quantified as follows: the size of macrophage ac-

cumulation core is 1.0 mm (lateral) by 0.79 mm (depth)

with a 36 μm thickness the fibrous cap, while the sizes

Y. X. Mao et al. / J. Biomedical Science and Engineering 4 (2011) 755-761

760

of the calcified regions are 0.6 mm by 0.4 mm (left side),

0.3 mm by 0.15 mm (right side).

To view changes of the clinical aspects of the plaque

along the direction of blood flow, a series of the

cross-section OCT images were taken from the positions

along the blood flow. Figure 5 shows six images of the

coronary from the WHHLMI rabbit in 0.1 mm apart

along the blood flow acquired utilizing the QSS-OCT

system with a 25 dB gain SOA inserted in sample arm.

Each image is in the same size as that shown in Figure 4.

These images formed a three-dimensional (3D) view of

the coronary of the WHHLMI rabbit. Size and density

changes of the clinical aspects of the coronary along the

blood flow were clearly distinguished from the 3D view.

A lipid core (black arrows), several calcified regions

(white arrows) with large size change in the 0.6 mm dis-

tance, and another lipid-rich area (grey arrows) beneath

the calcified regions. This lipid-rich area becomes larger

to a size of 1.5 × 1 mm2 along the calcified regions

shrink, while its fibrous cap (black diamond) remains

Figure 4. Ex vivo OCT images of a coronary from a WHHLMI rabbit acquired by the QSS-OCT system with the sample signal am-

plifier (a) and without (b). A-scan signals at the positions of the dashed arrow lines with pixels 100 (c), 500 (d), and 750 (e) acquired

by the QSS-OCT system with the sample signal amplifier (black) and without SOA (grey).

Figure 5. Ex vivo OCT images in another location of the coronary from the WHHLMI rabbit acquired by the QSS-OCT system with

the sample signal amplifier. Each image is taken from the positions in 0.1 mm apart.

JBiSE

Y. X. Mao et al. / J. Biomedical Science and Engineering 4 (2011) 755-761 761

around 0.25 mm.

5. CONCLUSION

High quality images of descending aorta harvested from

WHHLMI rabbits are produced by using a quadrature

swept-source OCT system containing a semiconductor

optical amplifier in the sample arm. A significant in-

crease in signal-to-noise ratio was obtained by inserting

an SOA in the sample arm of the QSS-OCT system. The

penetration depth of the QSS-OCT image was increased

with the addition of sample signal amplification. Pre-

liminary results show that vulnerable plaque with fibrous

caps, macrophage accumulations and calcifications pre-

sent in arterial tissue are measurable with our QSS-OCT

system. Our new QSS-OCT system reported in this work

could help in identify the locations of vulnerable coro-

nary plaques, in vivo, and in monitoring, with a high

degree of detail, the outcomes of coronary interventions.

6. ACKNOWLEDGMENTS

The authors are grateful to Dan P. Popescu, and Michael G. Sowa from

Institute for Biodiagnostics, National Research Council Canada, for

providing the coronary samples of WHHLMI rabbit.

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