Abstract

Ovarian cancer is fairly common, affecting one in every 70 to 100 women, and is the fifth

leading cause of cancer death among women. In the early stages of ovarian cancer, symptoms

are either not present or are mild and often disregarded. Current detection methods include

biopsy, a pelvic and rectal examination, transvaginal ultrasound, and a blood test for a tumor

marker, but problems exist with each. Optical spectroscopy, especially Raman, has shown great

promise recently for cancer detection. A Raman spectrum is a plot of scattered light intensity

versus the frequency shift of the scattered versus the incident photon. The peaks in the plot are

narrow and highly specific to a particular chemical bond, so each molecule has a unique

spectrum or “fingerprint” associated with it. This “fingerprint” allows differentiation between

normal and cancerous tissue. The goal of this project was to come up with an innovative Raman

probe design to allow reliable, minimally invasive detection of ovarian cancer. The in vivo

Raman probe should: fit through a microlaparoscope; be able to visualize the location of the

probe tip; read only the Raman signal of the tissue; be in direct contact with the tissue during

each measurement; and not induce a harmful reaction in the body. The probe was built

according to these specifications, and then tested by replacing the probe from Dr. Mahadevan-

Jansen’s clinical cervical system with our probe and following the same procedure as for the

cervical system. Although our prototype does not function up to the standards we set for it, it

gave us a great deal of hands-on experience in taking an idea all the way from conception to the

drawing board to the lab bench. The benefits that can be associated with this system are

thousands of saved lives, reduced health care costs, saved fertility, and reduction of major

surgeries. These benefits are important enough to have others continue this project and work to

refine it and develop later generations that can eventually be used clinically.

1

Introduction

Ovarian Cancer

Ovarian cancer is fairly common, affecting one in every 70 to 100 women, and is the fifth

leading cause of cancer death among women. In 2003, it is estimated that over 25,000 women

will be diagnosed with the disease, while over 14,000 women will die from it. The current five-

year survival rate for all cases is approximately 50 percent, but if it is detected before it has

spread beyond the ovaries, that number increases to around 90 percent. Family history plays a

large part in the disease as well; women with an afflicted primary relative have a 3.6 times

greater chance of developing the disease themselves [1].

In the early stages of ovarian cancer, symptoms are either not present or are mild and

often disregarded. One reason the disease is rarely detected until it has reached more harmful

stages is the minute size of the ovaries, as

demonstrated by Figure 1, which shows an

excised piece of cancerous ovarian tissue.

During stage one of ovarian cancer, the growth

of the tumor occurs solely in the ovaries, and

stage two occurs when the cancer develops a

pelvic extension. In stage three, the cancer has

spread beyond the ovaries, often to the lymph

nodes, and stage four sees the cancer spreading

outside the peritoneal cavity [3].

After the cancer reaches stage three and spreads outside of the ovary, the five-year

survival rate rapidly declines to approximately 15 to 20 percent; therefore, early diagnosis of

2

Figure 1 Picture of an excised ovarian tumor; note the size – units in mm [2].

ovarian cancer is essential for complete recovery. Unfortunately, the diagnosis is currently made

at stage three or later in 75 percent of cases [4]. This inability to detect the disease in its early

stages has severe economic impacts in addition to its harmful physiological impacts, as the

complications brought about by late-stage diagnosis cause the cost of treatment to increase by

approximately $40,000 over a patient’s lifetime [5].

Current Detection Methods

Current minimally invasive methods for detecting ovarian cancer include a pelvic and

rectal examination, transvaginal ultrasound, and a blood test for a tumor marker, but problems

exist with each. The pelvic and rectal examination does not detect ovarian cancer in its early

stages because of the size and location of the ovaries, and although ultrasound can detect tumors

in the ovary, it cannot determine whether they are cancerous. There have not been enough

clinical trials to support the efficacy of this method either. The tumor marker examined in the

blood test is the ovarian protein CA-125. Elevated levels of this protein are often an indication

of ovarian cancer, but increased levels of this factor are also associated with pregnancy, liver

disease, and endometriosis, giving it a positive predictive rate of only 10%. Thus, despite its

inherent invasiveness, biopsy remains the gold standard for diagnosing ovarian cancer since it is

the only method that provides the necessary accuracy and reliability for this task [1].

Optical Spectroscopy

When laser light is incident upon human tissue, it can be either scattered or absorbed by

that tissue. These light-tissue interactions provide a wide range of information about the tissue,

including the presence of certain pathologies, and given the rapid response time for these

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interactions, they can be measured essentially in real-time. Various methods of optical

spectroscopy that make use of these properties have been proven to be effective as real-time,

non-intrusive, automated diagnostic tools for several physiological systems, including the colon,

cervix, skin, and esophagus. More specifically, Raman spectroscopy has been shown to quickly

give accurate biochemical information necessary to diagnose cancer. A system developed by Dr.

Anita Mahadevan-Jansen that uses Raman spectroscopy to diagnose cervical cancer was recently

shown to have a 96 percent sensitivity and a 92 percent specificity in its diagnosis [6].

Raman Scattering

Raman spectroscopy utilizes its namesake inelastic scattering, which occurs when an

incident photon causes a scattering molecule to enter a virtual excited state, and then return to a

ground state either higher or lower than the

original through the emission of another photon.

In contrast to fluorescence, which involves

transitions between electronic energy levels,

Raman scattering exploits smaller transitions

between vibrational energy levels. Raman

Stokes scattering occurs when the scattered

photon has less energy than the incident photon,

which can be seen in Figure 2a, while Raman

Anti-Stokes scattering occurs when the scattered

photon has more energy than the incident

photon, as seen in Figure 2b [6].

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(a) (b)Figure 2: (a) Raman Stokes - incident photon has more energy than the scattered photon (b) Raman Anti-Stokes- incident photon has less energy than the scattered photon [6].

Figure 3 Raman spectra of normal and cancerous ovarian tissue

Raman Spectroscopy

A Raman spectrum is a plot of scattered light intensity versus the frequency shift of the

scattered versus the incident photon. The frequency shift is expressed in units of wavenumber,

which is the reciprocal of the wavelength, and thus proportional to frequency. The spectrum

consists of a series of peaks, each of which represents a different vibrational mode of the

scattering molecule. These peaks are narrow and highly specific to a particular chemical bond,

so each molecule has a unique spectrum or “fingerprint” associated with it. Thus, in a mass of

tissue, a specific interaction or

“fingerprint” can be isolated

from the overall spectrum in

order to determine the

chemical composition of the

tissue, which allows

differentiation between normal

and cancerous tissue. This can

be seen in Figure 3, which

shows the Raman spectra of normal and three kinds of cancerous ovarian tissue. Since nuclei

enlarge during cancer, it is hypothesized that certain protein and DNA peaks increase in intensity

in cancerous versus normal tissue. In Figure 3, slight increases in intensity can be seen at the

1330 cm-1 DNA peak and 1650 cm-1 protein peak, and a significant increase can be seen at the

1450 cm-1 protein peak. The Raman signal is very weak due to its dependence on small energy

transitions, so it requires a very strong source and sensitive detectors, as well as methods to filter

out signals from other types of interactions and the Raman signal of the fiber optics [6].

5

Goals

Our initial goals for this project were to gain a comprehensive understanding of Raman

spectroscopy and a basic understanding of ovarian physiology. We then sought to become

familiar with Dr. Mahadevan-Jansen’s Raman probe for cervical cancer, as we would be using

the same general system components, but designing a probe that could fit through a

microlaparoscope to reach the ovaries. We thus had to research the design limitations imposed

by the microlaparoscope and the physiological environment, and use this information to come up

with an innovative Raman probe design to allow reliable, minimally invasive detection of

ovarian cancer. The long-term goal of this project is to implement the design in a clinical

setting, although that was not an immediate goal for this semester.

Methodology

In accordance with the goals stated above, we did much reading and researching about

Raman spectroscopy and the latest developments in its use to detect cancer. We then

familiarized ourselves with Dr. Mahadevan-Jansen’s cervical Raman system, which is shown in

Figure 4. This system includes a diode

laser with a wavelength of 785 nm and 80

mW of power, a spectrograph, a charge-

coupled device (CCD) camera, a probe,

and a computer. The laser light travels

through the 400 μm core diameter fiber

optic into the probe, where a bandpass

filter eliminates any Raman signal from

the fiber and “cleans up” the laser light to

6

Figure 4: Basic Schematic of a Raman System.

ensure that only one wavelength is incident upon the sample. The transmitted light then interacts

with the sample, and back-scattered light reenters the probe tip, where it passes through a notch

filter that prevents the excitation laser light from entering the collection fibers and ensures that

only scattered light with a lower frequency can enter the bundle of collection fibers. The filtered

light then enters the spectrograph, at which point over 50% of the light is lost due to the size of

the fibers and the entrance slit of the spectrograph. The CCD camera then records the spectrum,

which is displayed on the computer.

Before measurements are taken, the camera must be cooled with liquid nitrogen to be at a

constant temperature of - 90˚C, and the parameters for data acquisition are set in order to ensure

that data is taken in the appropriate range. The probe is placed in direct contact with the sample,

which is necessary for collecting the Raman signal, and the laser and collection apparatus

operate for five seconds. After the tissue measurements, Naphtalene, 4-acetoamidophenol, and a

Neon:Argon light spectrum are all measured as standards to account for system variation. The

raw spectrum also includes fluorescence, which is somewhat reduced by taking measurements in

the dark, and noise, so it is necessary to process the raw spectrum to isolate the Raman

component. Fluorescence is removed using an automated polynomial fitting program based on

a fifth order polynomial, and noise is reduced using a Sawitsky-Godway filter that is

incorporated within the program. The final result is a Raman spectrum with a normalized mean

that can be used for data analysis.

With this complicated process in mind, we considered several possibilities for creating a

new probe that could use the same system (laser, spectrograph, and CCD) and fulfill certain

criteria. The in vivo Raman probe should: fit through a microlaparoscope, which has an internal

diameter of less than 3 mm; be able to visualize the location of the probe tip; read only the

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11 100μm fibers

Raman signal of the tissue; be in direct contact with the tissue during each measurement; and not

induce a harmful reaction in the body. The ideal solution, which would achieve maximum

efficiency for the signal collected, would have been to essentially minimize the existing cervical

probe with the filters at the probe tip. Due to the size constraints, however, the necessary parts

were either not available or prohibitively costly. With the help of an Innovation Work Bench

analysis [7] and discussions with our advisor and the

science machine shop, we settled on placing all of the

optics in a box outside of the probe itself, as shown in

Figure 5. We used ½ inch (12.7 mm) diameter optics

because that is the smallest size commonly available

from stock supplies, so that would keep both the box size

and costs down as much as possible. The lenses are

made of calcium fluoride, a material that does not

contribute to the Raman signal in our region of interest,

and have focal lengths chosen to keep the box size to a

minimum and to couple light from and into the fibers as

well as possible. Knowing the characteristics of

lenses available, and that the fibers (or bundles) should

have a core diameter of 400 μm since that is what the laser was designed for, we made a basic

AutoCAD drawing showing the placement of the optics relative to each other (Appendix A-1).

After further consultation with the machine shop, we made a more advanced series of AutoCAD

drawings (Appendices A-2 through A-4) to guide the production of the aluminum box.

8

As seen in the Appendices and Figure 5, the excitation leg (the shorter leg) contains a

bandpass filter that is centered at 785 nm with a bandwidth of 10 nm, and a biconvex lens with a

12.7 mm focal length to couple the light from the single laser fiber to the single probe excitation

fiber, each with 400 μm core diameters. The distances specified were derived from the basic

lens equation, which says that , where f is the lens focal length, i is the image

distance, and d is the object distance. These distances also work well since the numerical

aperture (NA) of all the fibers used is 0.22, so using the relation that NA = sin α for an air-fiber

interface, where α is the acceptance angle of the fiber, we should have light entering the fibers at

12.7o, and actually have an angle of 14.2o, meaning that little light is theoretically lost in the

coupling process.

On the collection side, each plano-convex lens has a 25.4 mm focal length, and is the

same distance away from their respective coupled fibers so that the beam is collimated when it

passes through the long-pass filter and then gets re-focused into the collection fiber bundle. The

LP filter has an optical density of 6 for wavelengths less than 790 nm, so no laser light or Raman

Anti-Stokes scattered light can pass through. The collection fiber bundle consists of 11 fibers

with a 100 μm core diameter to give the bundle an available area similar to that of the 400 μm

fibers, and then becomes a line of fibers on the other end since the spectrograph requires such an

arrangement. The fibers are screwed into the box via SMA connectors inserted into positioning

devices with two degrees of freedom – horizontal (x) and vertical (y) movement, that allow for

corrections necessary in coupling the light. As seen best in Appendix A-4, the optics are held in

place by small optic mounts, which are themselves mounted on a system of rails that allow

translation in the z-direction. Since lenses are not made with exact precision, this last degree of

freedom corrects for any errors in focal length to help couple and collimate the light, where

9

appropriate. Finally, the probe ends of the excitation and collection fibers were encased with

stainless steel tubing 2 mm in diameter that would be inserted into the microlaparoscope. This

tubing does not harm the body, and it allowed the two fibers to be placed right next to each other

to achieve maximum overlap in excitation and collection areas without using more advanced

techniques such as a ball lens or polishing the fibers at an angle.

To test the probe, we replaced the cervical probe with our probe and followed the same

procedure as outlined above for the cervical system. When it was clear that it was not working

as hoped, we took the optics out of the box and set up the system on the bench with the same

spacing as in the box. In this setup, we examined how much light was being transmitted through

each optical component, and tried to determine the reasons for losing substantially more light

than a theoretical analysis showed. After correcting any problems we found that could be

corrected with our given time and resources, we returned the optics to the box, tried to refine the

system in that state again, and then ran one final series of tests with our probe.

Results

Probe Data and Testing

The picture to the right in Figure 6 is

an image of the box described above, showing

the placement of the filters and lenses. The

lower part of the box is the excitation leg, and

the upper part is the collection leg. The

optical components were aligned to try to get

the maximum power transmission through

10

Figure 6: The box containing the optical components in their best arrangement.

each component, first on the bench top to

allow exact coupling and focusing distances

to be determined, and then within the box

itself. The metal rods within the box that

allow z-translation made ideal coupling

difficult because the placement is critical,

and the rods leave room for error.

Using the completed prototype seen

in Figure 7, the maximum power originally obtained from the laser through the entire excitation

leg was 15 mW out of 180 mW entering the system. Since 8.3% is not an ideal power

transmission, we took the optics back out of the box as described above, and determined the

losses that were occurring for each individual lens, the bandpass filter, the entire excitation leg,

the entire collection leg, and the fiber interfaces. A summary of the actual power transmitted

through each component and section can be seen below in Table 1, which also shows theoretical

transmissions according to Fresnel’s Laws and stated values from transmission curves.

Interface or system component Theoretical Power Transmission (%)

Actual Power Transmitted (%)

Biconvex lens 96 70Bandpass filter 60 55Entire excitation leg (biconvex lens and bandpass filter)

58 41

1st plano-convex lens 96 622nd plano-convex lens 96 90Both plano-convex lenses together 92 55Entire collection leg (with long pass filter estimate*)

80 48

Fiber-air interfaces (total of 8) 72 72Total 33 14

Table 1: The power transmission at various points for the Raman probe* We could not measure the transmission through the long pass filter since we did not have a laser of appropriate wavelength, so we are assuming its stated transmission value of 87%

11

Figure 7: The complete prototype.

As seen in the table, the maximum power that can be transmitted from the laser to the

spectrograph with this benchtop setup is only 14% of the original laser power, which is only 42%

of what could theoretically be transmitted. Since the laser has a maximum power of 300 mW,

the maximum power transmitted to the spectrograph is 42 mW. These calculations assume that

there is no loss of power in the signal compared to the excitation beam as well, which may or

may not be the case. In addition, when the optics were placed back in the box for the final time,

the transmission through only the excitation leg went back down to around 10%, which was still

an improvement over the original 8.3%.

One of the first spectra taken with our probe is shown below in Figure 8. There is a

Raman signal present in this graph, but it is not the signal that is expected; a spectrum with more

defined peaks at certain

wavenumbers should have been

produced. The alignment between

the spectrograph and the collection

fibers may not have been correct, as

we did not have the proper

connectors available, and this result

could also be noise from the

components of the Raman signal

from the probe itself. Longer

integration times making up for the large losses failed to significantly improve the images, so

more testing of the system needs to be done in the near future so that the system’s reliability can

reach that of the analogous cervical system, which has a 96% sensitivity and 92% specificity.

12

0

0.5

1

1.5

2

2.5

3

3.5

4

950 1150 1350 1550 1750

Raman Shift (cm-1)

Inte

nsity

%

Figure 8: One of the first Raman spectrum taken with the probe.

Economic Impact and Benefits

There are approximately 1,000 major medical hospitals and 20,000 gynecologists in the

United States. If the Raman system we have begun to develop were to be as prevalent as

mammography devices, there would be at least two machines, usually more, in each major

hospital. In the beginning, the Raman system would be used in conjunction with laparoscopic

procedures. The biggest problem with this is that in order for a laparoscopic procedure to be

done, some sort of abnormality must already exist. If this system were proved to be effective,

then the ovarian cancer screening would be a necessary part of every laparoscopic procedure,

which most gynecologists perform two or three times per week. Also, if this procedure is

available and not preformed, then the doctor could actually be sued for malpractice if the patient

is later diagnosed with ovarian cancer. The hope is that it could eventually become a screening

system that could be used on a more regular basis for women that have a family history of

ovarian cancer, in order to detect the cancer in its very early stages [8].

The major benefits of the system include saving the fertility of women who would

otherwise have their ovaries removed unnecessarily, decreasing the need for some major surgery,

and once it is used as regular screening system, saving thousands of lives per year. It has the

potential to increase the total five-year survival rate for this disease by up to 40 %, and thus save

5000 lives each year, by virtue of the marked increase in survival rate with early detection [1]

The cost of the entire system minus the computer would be about $15,800, and the

breakdown of this price can be seen in Table 2. The list shown below does not include the initial

design price either. The student work price for the initial design and testing would be $4500,

from 2 students working for 150 hours @ $15 per hour. With a slight increase in price, of

approximately $1,000, the student work time could be greatly decreased because instead of

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Component Rate PriceProbe Replace yearly $2100CCD camera One time

$12,000Spectrograph One timeLaser One timeStudent Work 45 hours @ $15 per hour per student $1350Machine Shop Work 10 hours @ $35 per hour $350Total $15,800

Table 2: The price breakdown for the Raman system.

assembling parts, they could order them pre-assembled. The hospital would also have to possess

a laparoscope or a microlaparoscope, but all hospitals should already have this equipment. The

probe itself would probably have to be replaced on a yearly basis to maintain integrity, but the

rest of the system would not require a great deal of up-keep, needing only to be replaced with

significant advances in technology or old age.

A system that has the potential to detect ovarian cancer is such a desired product in the

medical field that the overhead for marketing this device would be minimal. The best ways to

accomplish this would be word of mouth and articles and ads in the appropriate journals. There

are several obstetrics and gynecology conferences every year, so a presentation at a large

conference would allow the doctors to become informed and the device to come into market.

In laparoscopic surgery, the complication rate is less than 1 in every 200 women, and

microlaparoscopy is known as the “non Band-Aid” surgery since the incision required is so small

that it does not require stitches or leave a scar [9]. Since this device would be used in

laparoscopic surgeries, the complication rate is not expected to be very significant. Although

laser light is used, the wavelength, intensity and exposure time are not going to damage the skin

or internal organs; however, it is still important to be careful when working with lasers that the

laser light is not directly shone into anyone’s eyes. There are several long cords containing the

fiber optics that need to be considered when working with probe. Although entanglement is not

14

a significant problem, the surgeon and the nurses still need to be cautious of the wires. Other

potential safety hazards identified by a designsafe analysis [7] include contact with biological

tissues, problems from posture during the procedure, and knocking the system components off

the cart. The analysis also showed that with proper precautions, none of the above safety

concerns should have a major impact on the use of this device, and any possibility of slight risks

are far outweighed by the potential benefits.

Another concern for a device like this is sterilization. There are several methods for

sterilizing medical instruments, some of which involve heat, and other others that involve being

sprayed with a gas. The best method of sterilization for the Raman probe is gas since the fiber

optics cannot be exposed to a tremendous amount of heat. As long as the device is properly

sterilized and biocompatible materials, such as a sapphire window, are used for the probe tip, the

FDA should not have many issues with approving it because of its similarity to the cervical

probe in the process of approval right now, and because it uses a procedure (laparoscopy) that is

well established in the medical community.

Conclusions

Although our prototype does not function up to the standards we set for it, it gave us a

great deal of hands-on experience in taking an idea all the way from conception to the drawing

board to the lab bench. It has laid the foundation for future work on this device by identifying

problematic areas and by demonstrating that the basic design shows promise. We succeeded in

making a probe that will fit through a microlaparoscope and take a Raman spectrum, but issues

related to light coupling and transmission prevented it from working to its full capacity.

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Specifically, unexplained losses occurred in both the biconvex lens and the first plano-

convex lens as shown previously in Table 1. These may be due to the hygroscopic nature of

calcium fluoride, some damage that was inflicted upon them, a possible flaw in the anti-

reflective coatings, or something else all together. Since this was our first time working with

extremely small fiber optics, our ability to make fiber connections, bundles, and lines was not

exactly up to par with commercial companies, so that contributed to greater losses as well.

Because the distance between fibers on opposite ends of the box could not be varied, we had to

assume that the focal length of the biconvex lens was exactly as stated in making the length

calculation. This assumption was probably not correct, however, so we could not achieve ideal

coupling in the excitation leg. The greater losses when the optics are placed in the box are most

likely explained by increased reflectance off the metallic surfaces, as well as less freedom in

positioning the components as compared to doing so on the bench. Finally, the placement of the

optics away from the probe tip may be creating too much Raman signal from the fiber itself.

Recommendations

We have a number of recommendations that could make this device, or a later generation

of it, perform at a higher level. To eliminate the problem of too much reflectance, the box

should be painted with a flat black paint, and the lenses need to be thoroughly examined to

determine why two of them are reflecting so much light. They would probably have to be

replaced, perhaps with higher quality ones with a better anti-reflective coating. By ordering

professionally made fibers, especially the bundle-to-line one, we could achieve more efficient

coupling and transmission of light. By replacing the biconvex lens in the excitation leg with two

plano-convex lenses in a manner similar to the collection leg, we would eliminate the problem of

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having improper and static coupling distances, and would possibly achieve greater transmission

through the bandpass filter as well since a collimated beam will pass through it more easily than

a diverging beam will. This would also make the box more uniform in shape. If we could make

the components translate in the z-direction more smoothly, we would have an easier time

coupling the light, and it would help to have the components more stable as well, so that they do

not get out of alignment while being moved. Some degree of rotational freedom may also

increase throughput of light.

These changes could be implemented in the current prototype, but progressing to a

second generation model may be the way to go. The box could probably be made less bulky, and

perhaps not even be the shape of a traditional box, but rather be a tube. At the tip of the probe, a

ball lens should be attached, or the fibers polished at an angle, to increase the spot size covered

by the light coming out of the probe. To protect the tips of the fibers or ball lens, a sapphire

window could be added as well. Other potential improvements for future generations include

developing a probe that could look past the surface of the ovary, and taking multiple images in

rapid succession.

For whatever design is ultimately chosen as the one to pursue deeply, many statistical

comparisons will have to be made with previously proven probes to determine the accuracy and

precision of the instrument. Once the in vitro tests have confirmed its efficacy, the probe should

then be entered into clinical trials, following all of the regulations of the FDA and the

Institutional Review Board. Although we will not oversee this process, we hope that the

foundations we have laid with this work will be continued by others in the same lab, and will

eventually lead to many saved lives.

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Acknowledgments

Dr. Anita Mahadevan-Jansen – our primary advisorMark Mackanos, Chad Lieber, Amy Robichaux – graduate studentsDr. Gary Kanter – OB/GYNDr. Paul King – course instructor

References

1. National Cancer Institute: http://www.nci.nih.gov/cancer_information/cancer_type/ovarian/

2. http://www.pathguy.com/bryanlee/ov_ca1.jpg

3. http://www.ovarian.org/

4. http://www.ovariancancer.org/content/1-5-1.html

5. http://www.hcfinance.com/dec/dectside2.html

6. Mahadevan-Jansen, A., Raman Spectroscopy: From Bench top to Bedside. (2003).

7. http://vubme.vuse.vanderbilt.edu/srdesign/2002/group9/

8. Kanter, Gary. Personal conversation. 2/18/03.

9. http://www.obgyn.net/hysteroscopy/gynetrends/microlaparoscopy.htm

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