Ultrasound Imaging of Bone for Spinal Fusion Surgery Guidance: Simulation

and Experimental Results

by

Al-Hassan Aly

A thesis submitted in conformity with the requirements for the Degree of Master of Applied Science

The Institute of Biomaterials and Biomedical Engineering

University of Toronto

© Copyright by Al-Hassan Aly (2010)

ii

Abstract Ultrasound Imaging of Bone for Spinal Fusion Surgery Guidance: Simulation and

Experimental Results Al-Hassan Aly

Master of Applied Science, 2010 The Institute of Biomaterials and Biomedical Engineering

University of Toronto

In order to continue development of an ultrasound-guidance system for spinal

fusion surgery, simulation and experimental research was conducted to study the effects

of bone on ultrasound imaging. Simulation work examined the effect of bone volume and

transducer frequency on image quality and accuracy. Experimental work utilized a

3.2MHz prototype ultrasound probe to create ultrasound images of pedicles. The

simulation results, based on an idealized anatomical model, provided higher-quality

images than the experimental results. It was determined that high bone volume and high

transducer frequency have a detrimental effect on image quality. The experimental results

suggest that the high variability in pedicle shape results in variability in ultrasound image

quality. Overall, the simulation and experimental results suggest that ultrasound imaging

of bone is feasible at relatively low frequencies, while highlighting the need for more

experiments to take into account the substantial variability in pedicle shape and bone

volume.

iii

Acknowledgments I would sincerely like to thank Professor Richard S.C. Cobbold and Dr. Howard J.

Ginsberg for their guidance, advice, and support for this research. Professor Cobbold’s

teaching on the physics of ultrasound and ultrasound systems has been invaluable in

conducting this research, while Dr. Ginsberg’s clinical experience helped me see the “big

picture” of how this research relates to existing health care protocols for spinal fusion

surgery. I would also like to thank Dr. Ginsberg for the opportunity to visit the operating

room and observe this procedure. Both Professor Cobbold and Dr. Ginsberg have also

been extremely supportive in editing and preparing numerous abstracts, posters, and a

manuscript, as well as funding trips to conferences for presentations.

I would also like to thank Muris Mujagic, Charles Lai, Derek Wright, Renee

Warriner, Alexia Giannoula, and Hassan Masoom for their help and support. To the

Davies Lab I am especially indebted for their help with the µCT machine. Finally, the

staff at the Anatomy Department has been very helpful in the procurement of the

necessary vertebrae for this research.

In addition, we are grateful to the Natural Sciences and Engineering Research

Council of Canada for a supporting research grant.

iv

Table of Contents

Abstract .............................................................................................................................. ii Acknowledgments............................................................................................................. iii Table of Contents ............................................................................................................. iv List of Tables...................................................................................................................... v List of Figures .................................................................................................................... v Chapter 1 – Introduction .................................................................................................. 1

1.1 Anatomy of the spine .............................................................................................. 1 1.2 Spinal fusion surgery ............................................................................................... 9 1.3 Existing image guidance technologies for pedicle screw insertion....................... 18 1.4 Developing an ultrasound alternative .................................................................... 22

1.4.1 Acoustic properties of trabecular and cortical bone .................................. 22 1.4.2 Design of prototype transducer ................................................................. 27

1.5 Hypothesis ............................................................................................................. 29 1.6 Research objectives ............................................................................................... 29

Chapter 2 – Simulation Methods and Results .............................................................. 30 2.1 Creating a simulation model ................................................................................. 31 2.2 Simulation analysis and results ............................................................................. 42

2.2.1 Image accuracy .......................................................................................... 44 2.2.2 Attenuation ................................................................................................ 49

Chapter 3 – Laboratory Experiments ........................................................................... 52 3.1 Experimental protocol ........................................................................................... 53 3.2 Analysis of ultrasound images .............................................................................. 56

Chapter 4 – Summary and Conclusions ........................................................................ 66 4.1 Summary of contributions ..................................................................................... 66 4.2 Limitations ............................................................................................................ 68 4.3 Conclusions ........................................................................................................... 70 4.4 Future work ........................................................................................................... 70

References ........................................................................................................................ 71

v

List of Tables Table 2.1: Bone volumes for models and individual slices ............................................... 35 Table 2.2: Material properties .......................................................................................... 37 Table 2.3: Simulation parameters ..................................................................................... 41

List of Figures

Figure 1.1: Lateral view of the spine................................................................................... 2 Figure 1.2: Connections between adjacent vertebrae .......................................................... 2 Figure 1.3: Cervical, thoracic, and lumbar vertebrae ......................................................... 3 Figure 1.4: Thoracic and lumbar pedicle dimensions ......................................................... 6 Figure 1.5: Blood vessels of the spine ................................................................................ 7 Figure 1.6: Neuroanatomy of the spine ............................................................................... 8 Figure 1.7: Three-column method of fracture classification ............................................. 10 Figure 1.8: Example of lower thoracic burst fracture ....................................................... 10 Figure 1.9: Example of lower thoracic primary malignant tumour ................................... 11 Figure 1.10: Degenerative disk disease ............................................................................. 12 Figure 1.11: Placement of bone graft in spinal fusion surgery ......................................... 13 Figure 1.12: Surgical approaches to the lumbar spine ...................................................... 14 Figure 1.13: Post-operative posterior spinal fusion images .............................................. 17 Figure 1.14: Example of intra-operative fluoroscopy imaging ......................................... 19 Figure 1.15: Example of intra-operative 3D fluoroscopy imaging ................................... 20 Figure 1.16: Example of 3D infrared-based surgical navigation ...................................... 21 Figure 1.17: Correlation between bone volume and S.O.S. in trabecular bone ................ 24 Figure 1.18: Fast and slow waves in trabecular bone ....................................................... 24 Figure 1.19: Frequency-dependent attenuation in trabecular bone ................................... 26 Figure 1.20: Dispersion in trabecular bone ....................................................................... 27 Figure 1.21: Prototype ultrasound probe ........................................................................... 28 Figure 1.22: Use of the prototype probe ........................................................................... 28 Figure 2.1: Intra-pedicular imaging using ultrasound ....................................................... 32 Figure 2.2: Dilation and erosion of trabecular bone .......................................................... 34 Figure 2.3: General simulation geometry .......................................................................... 36 Figure 2.4: Material properties .......................................................................................... 38 Figure 2.5: Transmitted plane wave .................................................................................. 40 Figure 2.6: A-line simulation ultrasound images .............................................................. 43 Figure 2.7: Error vs. bone volume and frequency ............................................................. 45 Figure 2.8: Measured distance vs. bone volume and frequency ....................................... 45 Figure 2.9: Simulation A-mode images ........................................................................... 47 Figure 2.10: Attenuation dependence on bone volume and transducer frequency ........... 50 Figure 3.1: Pedicle and three-point needle holder ............................................................. 54 Figure 3.2: Experimental equipment ................................................................................. 55 Figure 3.3: Experimental A-line ultrasound images ......................................................... 57

vi

Figure 3.4: B-mode ultrasound images of left pedicle of L4 vertebra .............................. 59 Figure 3.5: Micro-CT and MIP images of left pedicle of L4 vertebra .............................. 60 Figure 3.6: B-mode ultrasound image and photographs of right pedicle of L4 vertebra .. 63 Figure 3.7: B-mode ultrasound images of right pedicle of L1 vertebra ............................ 64

1

Chapter 1 – Introduction

Over the last several decades, ultrasound imaging has become ubiquitous in

several fields of medicine and surgery. Abdominal, vascular, and emergency surgery as

well as obstetrics/gynaecology and oncology have benefited greatly from the introduction

of ultrasound imaging devices designed for those specialities. The wide range of uses of

ultrasound imaging requires ultrasound transducers to be designed specifically for their

intended application. In the process of designing an ultrasound system for a specific

surgical procedure, the acoustic properties of the specific tissues encountered in the

operating area must be taken into account. For the surgical procedure of interest in our

research – spinal fusion surgery – unique challenges exist due to the acoustic properties

of trabecular and cortical bone. Our goal is to provide the surgeon with a means to

accurately insert pedicle screws during spinal fusion surgery, thus minimizing

complications. To meet this challenge, research utilizing both computer simulations as

well as laboratory experiments has been conducted to study ultrasound imaging of bone.

First, it is necessary to examine the anatomy of the spine.

1.1 Anatomy of the spine

The main functions of the spine are to provide stability, protection of neurological

systems, and connections for muscles. The spine consists of 33 vertebrae: 7 cervical, 12

thoracic, 5 lumbar, 5 sacral and 4 coccygeal. Figure 1.1 shows a lateral view of the spinal

column. The first 24 vertebrae, from the first cervical vertebra (C1) to the last lumbar

vertebra (L5), are separate and mobile. The main connections between adjacent mobile

vertebrae are the articular facets and intervertebral disks (Figure 1.2).

2

Figure 1.1- Lateral view of the spine. Normal spine curvature consists of the primary kyphotic curves of the thoracic and sacral sections and the lordotic curves of the cervical and lumbar spine, which develop at a later age. The cervical spine supports the head and neck. The thoracic spine is connected to the ribcage through several joints. The sacral and coccygeal vertebrae are normally fused and immobile. Modified from Vaccaro et al. (2005).

Figure 1.2- Connections between adjacent mobile vertebrae. The intervertebral disk consists mostly of Type I and Type II collagen. The superior articular facet of the vertebra below articulates with the inferior articular facet of the vertebra above. The intervertebral disk and vertebral bodies bear approximately 80% of the load carried by

3

the spine, while the articular facets carry approximately 20%. Reproduced with permission from Vaccaro et al. (2005).

Important features present in all mobile vertebrae of the spine include the

pedicles, the transverse processes, the lamina, and (with the exception of C1) the

vertebral body and spinous process. However, the vertebrae of the different regions of

the spine exhibit vastly different morphologies. Since the vertebral bodies are the main

load-bearing areas, their size increases towards the bottom of the vertebral column (see

Figure 1.3). The shape of the spinous process and transverse processes (which provide

connections for ligaments and muscles) as well as the shape and size of the pedicles

varies between the cervical, thoracic and lumbar vertebrae. The cervical vertebrae have

the smallest vertebral bodies, due to the fact that they bear the least weight. The spinal

canal, containing the spinal cord, is formed by the anterior walls of the lamina, the

posterior wall of the vertebral body, and the medial walls of the pedicles. The spinal

canal is largest in the cervical vertebrae.

4

Figure 1.3- Cervical, thoracic, and lumbar vertebrae. The C1 (or atlas) vertebra has no vertebral body or spinous process. It instead joins the C2 (axis) vertebra via the dens and atlanto-axial articulation. The dens articulates with the anterior arch of C1. Ligaments hold the dens to the anterior arch of C1, thus allowing C1 to rotate about the dens. The transverse processes of the cervical vertebrae contain a small foramen (the transverse foramen) which encloses the vertebral artery. A large spinal canal, small vertebral body and narrow pedicles are characteristic of the cervical vertebrae. Reproduced with permission from Vaccaro et al. (2005).

5

Figure 1.3 (continued)- Thoracic and lumbar vertebrae are larger than cervical vertebrae: the transverse processes are longer and angled posteriorly, the spinal canal is smaller, and the lamina, pedicles, and vertebral body, which comprise the boundary around the spinal canal, are larger. The spinous process is longer and is angled caudally in the thoracic vertebrae. A thoracic vertebra articulates with the ribs through the superior and inferior costal facets and the transverse costal facets. In thoracic and lumbar vertebrae, the intervertebral foramen is bordered above and below by the pedicles of the adjacent vertebrae. Reproduced with permission from Vaccaro et al. (2005).

The pedicles are especially important in spinal fusion surgery. Roughly

cylindrical in cross section, they serve to connect the vertebral body to the lamina and the

spinous and transverse processes. The pedicles generally become larger in the lower

spinal column. This is due to an increase in both height and width. In the lumbar spine,

pedicle width and height are generally larger than in the thoracic spine (as shown in

Figure 1.4). It is also important to note that the width of the pedicle varies along the

length of the pedicle, and may be narrowest in the centre and wider where it meets the

vertebral body and the lamina. Furthermore, the angle that the pedicles make with the

vertebral body varies among the different regions of the spine; in the lower thoracic

6

spine, the angle is often almost perpendicular to the vertebral bodies, while in the lumbar

spine and the upper thoracic spine the pedicles are angled medially.

Figure 1.4- Thoracic and lumbar pedicle dimensions. Most vertebrae in the thoracic and lumbar spine have pedicles that are wider in the sagittal plane than in the axial plane, resulting in tall, narrow pedicles. The pedicle width increases towards the bottom of the spine to accommodate the weight of the thorax and abdomen. Reproduced with permission from Zindrick et al. (1987).

Besides the vertebrae, which consist of cortical bone, trabecular bone and bone

marrow, the spine also includes many soft tissues, blood vessels, and nerves. Several

ligaments are aligned in the cranio-caudal direction connecting the vertebral bodies and

posterior arches of adjacent vertebrae. Major blood vessels, including the aorta and vena

7

cava, intercostal arteries (in the thoracic spine), and radicular arteries are located around

the vertebrae and feed both the bony spine as well as the spinal cord, as shown in Figure

1.5. Additionally, in the cervical vertebrae, the vertebral artery travels within the

transverse foramen located just lateral to the pedicles.

Figure 1.5- Blood vessels of the spine. The aorta is located just left of the medial plane and anterior to the vertebral body. In the lumbar spine, the lumbar artery originates from the aorta and travels to the pedicles. Passing through the intervertebral foramen, it splits into anterior and posterior vertebral canal branches; the anterior branch feeds the vertebral body through the nutrient arteries, while the posterior branch supplies the posterior arch and the nerves of the spinal canal. Venous drainage is achieved through the intervertebral and lumbar veins located lateral and caudal to the pedicles. Reproduced with permission from Moore and Dalley (2006).

8

The spinal cord carries both sensory signals through the spinocerebellar and

spinothalamic tracts as well as motor signals in the corticospinal tract. Originating in the

foramen magnum above the C1 vertebra, it travels through the spinal canal to the first or

second lumbar vertebra, where it becomes the cauda equina. Nerve roots exit the spinal

cord at each vertebra in the spine, passing through the intervertebral foramen as shown in

Figure 1.6. The intervertebral foramen is bordered by the pedicles of the vertebrae above

and below the exiting nerve root; the nerve roots enervate the limbs as well as the visceral

organs such as the heart.

Figure 1.6- Neuroanatomy of the spine. The lumbosacral enlargement (shown as a bulging of the spinal cord from T11 to L2) contains nerve roots that continue in the lumbar spine as the cauda equina before exiting the spine. Except in the cervical spine, the nerve roots are named after the vertebra directly above. The spinal cord is covered by the pia, arachnoid, and dura mater (which collectively form the meninges). Close to the intervertebral foramen, the nerve roots are also enclosed by a dural root sheath. Reproduced with permission from Vaccaro et al. (2005) [left figure], and Moore and Dalley (2006) [right figure].

9

Knowledge of the vascular- and neuro- anatomy of the spine is essential to

minimize the risk of inadvertent damage to important blood vessels or nerves during an

operation. The close proximity of the spinal cord, nerve roots, and major blood vessels to

the pedicles poses the primary challenge in spinal fusion surgery.

1.2 Spinal fusion surgery

As mentioned previously, two of the major functions of the spine are to maintain

stability and protect the spinal cord. Many factors can contribute to a reduction in the

spine’s ability to fulfill those functions.

Trauma can result in fracturing of one or more vertebrae. Fractures are

categorized according to the location and type of injury to the affected vertebra(e). Each

vertebra is divided into three columns, as shown in Figure 1.7. A particular injury can

affect one, two, or three columns in a single vertebra, and may affect several adjacent or

non-adjacent vertebrae. The type of injury to each column can be a compression, tension,

shear, rotation, or burst fracture. For example, some types of fractures may consist of

compression of the anterior column and tension of the posterior column, or tension of the

anterior column and compression of the posterior column. A fracture can cause a

dislocation of part of the bony spine and a subsequent compression of nearby nerve roots

or the spinal cord. This can cause permanent damage to the affected nerves due to

necrosis (cell death), inflammatory response and the death of myelin cells (Onose et al.

2009). Figure 1.8 shows an anterior column compression fracture with a burst fracture of

the middle column, causing a compression of the thecal sac (the membrane surrounding

the spinal cord). Surgical indications generally depend on many factors including the

10

degree of neurological damage, amount of excessive kyphosis or lordosis, and current

stability of the vertebral column.

Figure 1.7- Three-column method of fracture classification. The posterior column includes the spinous process, transverse processes, superior and inferior articular facets, lamina, and pedicles. The middle column consists of the posterior half of the vertebral body, and the anterior column consists of the anterior half of the vertebral body. The interspinous ligament (which connects adjacent spinous processes) is also shown; this ligament and other parallel ligaments connecting adjacent vertebrae can be torn in some types of fractures involving the posterior or middle column. Reproduced with permission from Vaccaro et al. (2005).

Figure 1.8- Example of lower thoracic burst fracture. This MRI shows the posterior wall of the vertebral body (middle column) impinging on the spinal cord, with compression of the anterior column. The vertebra affected is the T12 vertebra. Reproduced with permission from Vaccaro et al. (2005).

11

Primary and metastatic tumours of the spine can lead to neurological damage or

spine instability. Primary tumours are differentiated based on whether they are benign or

malignant. Benign primary tumours, such as giant cell tumours and osteoblastomas, due

to their proximity to the spine, can prove fatal if not diagnosed and surgically resected or

treated. Metastatic tumours of the spine often originate from breast, prostate, or other

tissue carcinomas. A chondrosarcoma (a malignant tumour derived from cartilage cells)

is shown in Figure 1.9; originating from the costovertebral junction (the joint connecting

the vertebral body and the rib), the tumour has spread into the vertebral body and the

ipsilateral transverse process. Tumours of the spine can also apply pressure to nerve roots

or the spinal cord, causing symptoms of pain. Surgical resection of a tumour that has

invaded the bony spine may significantly diminish the stability of the spine and warrant a

fusion of the affected vertebra(e) with the adjacent vertebrae.

Figure 1.9- MRI of primary malignant tumour. The tumour boundary is shown by the white arrows. The vertebra shown is the T11 vertebra. Reproduced with permission from Vaccaro et al. (2005).

12

Degenerative disk disease is a common cause of pain and other neurological

symptoms. A loss of fluid in the intervertebral disk between vertebrae causes a reduction

in the height of the disk (Figure 1.10). The loss of disk height causes stenosis (narrowing)

of the spinal canal or the intervertebral foramen, causing compression of either the spinal

cord or the nerve roots, respectively. Surgical treatment may be warranted depending on

the level of the stenosis, the number of affected vertebrae, and the stability of the spine.

Degenerative spondylolisthesis (a forward displacement of a vertebra over the vertebra

below it) or scoliosis secondary to degenerative disk disease also indicates that fusion

may be necessary.

Figure 1.10- Degenerative disk disease. The illustration shows a lateral (top), axial (centre) and posterior (bottom) view of a normal vertebra on the left, and a vertebra affected by degenerative disk disease on the right. The collapse of disk height causes the intervertebral foramen to become narrower, resulting in compression of the exiting nerve roots. The expansion of the degenerated disk into the spinal canal applies pressure to the spinal cord. Surgical treatment in this case would seek to restore the distance between the vertebrae in order to reduce pressure on the nerves. Reproduced with permission from Vaccaro et al. (2005).

13

In most cases, surgical treatment consists of a spinal fusion procedure to eliminate

the pressure on the affected nerves (decompression) and restore stability (fusion). Fusion

consists of the immobilization of two or more adjacent vertebrae by arthrodesis (the

surgical fixation of a joint). Arthrodesis is accomplished by the introduction of a bone

graft in between the vertebrae; over the course of several months, the bone graft will

grow and fuse with the vertebrae, creating a bony bridge between them. The bone graft

may be placed either in between the vertebral bodies or across the transverse processes,

as shown in Figure 1.11. The bone graft is generally either autograft (obtained from a

different bone of the same patient) or allograft (obtained from a different human donor).

In addition to the bone graft, instrumentation (metal implants such as screws and rods)

may be utilized to immobilize the vertebrae and provide stability until the bone graft has

completed the fusion. The instrumentation is usually not removed after successful fusion

has occurred, due to the expense and risk of complications of a second procedure.

Figure 1.11- Placement of bone graft. In cases where the intervertebral disk has collapsed and caused narrowing of the intervertebral foramen, the degenerated disk may be removed and replaced by a bone graft (as shown on the left) which restores the width of the foramen and decompresses the nerve roots. The bone graft fuses with the vertebral bodies of the vertebrae. The bone graft is inserted under compression, that is, with the vertebral bodies compressing the graft from both sides; this reduces movement of the bone graft post-operatively and increases the likelihood of successful fusion. Alternatively, the vertebrae may be fused by the addition of small pieces of bone along

14

the transverse processes of the vertebrae, as shown on the right. The transverse processes (and in the case of lower lumbar fusion the sacral ala, located just lateral to the vertebral body of S1) can be decorticated (through removal of the cortical bone) to improve fusion with the bone graft. Modified from Vaccaro et al. (2005).

The surgical approach depends on the region of the spine that is affected

(cervical, thoracic, or lumbar) and the type of specific procedure that is planned. For the

lumbar spine, retroperitoneal and posterior approaches (Figure 1.12) allow for exposure

of the spine, decompression and fusion. For the thoracic spine, posterolateral and

transthoracic, and for the cervical spine, transoral, retropharyngeal and posterior

approaches can be used. The type of approach is also determined by whether

instrumentation will be implanted. Anterior approaches can be used to implant interbody

bone grafts and vertebral body screws, while posterior approaches allow for the insertion

of pedicle screws and postero-lateral fusion along the transverse processes.

Figure 1.12- The posterior approach. Retraction of the adipose tissue and muscles allows visualization of the spinous process, lamina, and transverse process. The posterior approach allows decompression of the posterior aspect of the vertebra, relieving pressure on the nerve roots. The intervertebral disk can be excised, if necessary, through a window

15

created in the lamina. The anterior aspect of the vertebra can also be indirectly decompressed using the posterior approach. Finally, the insertion of pedicle screws and placement of the bone graft can be accomplished with this approach. Reproduced with permission from Vaccaro et al. (2005).

Our research was conducted specifically for the posterior approach employing

pedicle screw fixation. The insertion of the pedicle screws is particularly challenging and

is greatly aided by surgical guidance technology.

The first step in spinal fusion surgery is the exposure of the spine. Using the

selected approach, an incision is made through the skin, and the soft tissue surrounding

the spine is cleared using a scalpel or electrocautery. The muscle and adipose tissue is

retracted to provide a clear view of the spinous process and the transverse processes (in a

posterior approach). A laminectomy may be performed to allow access to the

intervertebral space; this allows the intervertebral disk to be excised and the bone graft to

be inserted if an interbody graft is being used.

Before a pedicle screw can be implanted, the entry point must be identified in the

posterior arch. A cannulation probe is then advanced by hand through the cortical bone

and into the trabecular bone, after a small entry hole is created with a cortical bone burr.

As the probe is advanced into the pedicle, tactile feedback provides the surgeon with vital

information about the type of bone (trabecular or cortical) the probe is disrupting. A

sudden increase in the resistance of the bone to the cannulation probe could be due to the

fact that the cannulation probe is in contact with cortical bone. On the other hand, an

abrupt drop in bone resistance may be due to the cannulation probe having been advanced

through the cortical bone into the soft tissue. Such a breach of the cortical bone can lead

to serious complications if the cannulation probe punctures a blood vessel or lacerates a

16

nerve. As the anatomy of the spine indicates, critical vessels and nerves are located

around the pedicle, especially in the medial, superior, and inferior directions.

The bore hole created by the cannulation probe provides guidance for the

subsequent insertion of the pedicle screw. The pedicle screw diameter is generally

selected to be the largest that can be accommodated by the pedicle; the diameter and

angle of the pedicle is obtained from pre-operative images. The insertion of the pedicle

screw carries the same risk as the cannulation of the initial bore hole; breaching of the

cortical wall of the pedicle is possible, especially if multiple bore holes were created in a

process of trial and error. The grip of the pedicle screw threads on the trabecular bone, in

combination with the rigidity of the rods placed across the screws, completes the fixation

of the vertebrae. As shown in Figure 1.13, the pedicle screws extend well into the

vertebral body, are lateral to the spinal cord, and superior to the exiting nerve root. These

images were acquired with post-operative medical imaging technology. Intra-operative

technology is also available and being developed to provide surgeons with information

about the placement of the pedicle screws during surgery, when mistakes can be more

easily corrected.

17

Figure 1.13- Post-operative posterior spinal fusion images. Top left: Post-operative axial reconstruction from a CT scan following a thoracic spinal fusion. Top right: Schematic view of pedicle screw placement, showing the relative positions of the screws, the spinal canal, and the nerve roots. Bottom left: Post-operative lateral X-ray of a lower lumbar postero-lateral fusion (fusion along the transverse processes) with pedicle screw and rod instrumentation. Bottom right: Post-operative lateral X-ray of interbody fusion with pedicle screw and rod instrumentation. Follow-up imaging is also conducted to verify that successful fusion has occurred between the vertebrae by observing the growth of new bone (appearing as a solid white mass) either in the intervertebral disc space or along the transverse processes. Reproduced with permission from Hartl et al. (2004) [top left figure], Vaccaro et al. (2005) [top right and bottom left figures], and Mandigo et al. (2007) [bottom right figure].

18

1.3 Existing image guidance technologies for pedicle screw insertion

While remaining a primary source of information, tactile feedback is limited by

the fact that the resistance of the bone to cannulation can vary considerably depending on

the patient’s age and bone volume. Conditions such as osteoporosis can also result in a

very low level of bone resistance to cannulation, reducing the effectiveness of tactile

feedback in differentiating between bone and soft tissue.

Therefore, several medical imaging modalities have been employed in spinal

fusion surgery to assist the surgeon. Fluoroscopy images provide a two-dimensional

projection, usually a lateral projection, of the spine (as shown in Figure 1.14).

Fluoroscopy images taken during cannulation of the pedicle or placement of the pedicle

screw can provide the surgeon with important information about the accuracy of the

approach in the sagittal plane, i.e. in the cranio-caudal directions. However, lateral

fluoroscopy images do not provide sufficient information on the accuracy of the pedicle

screw placement in the axial plane (such as screw proximity to the spinal cord).

Moreover, fluoroscopy exposes the patient and the surgical team to ionizing radiation; the

total radiation exposure varies depending on the number of images acquired during a

given procedure. Fu et al. (2008) reported a correct pedicle screw insertion rate using

fluoroscopy of 93.2%. On the other hand, Kotil et al. (2008) reported an incorrect pedicle

screw insertion rate of only 5.6% without using any fluoroscopic guidance.

19

Figure 1.14- Example of intra-operative fluoroscopy imaging. The top image shows a posterior projection of the spine; this provides a cross-sectional image of the pedicles. The bottom image shows a lateral projection, which provides a sagittal view of the pedicles. These fluoroscopy images are aided by navigation technology that allows a “phantom” image of a surgical instrument, such as a cannulation probe, to be superimposed on the fluoroscopy image, enabling the surgeons to ensure that the probe remains within the confines of the pedicle and vertebral body. Reproduced with permission from Kim et al. (2008) [top figure], and Fu et al. (2008) [bottom figure].

20

The first limitation of fluoroscopy is resolved through the use of three-

dimensional imaging. Three-dimensional fluoroscopy utilizes several consecutive

fluoroscopy images, taken from different angles, to reconstruct two-dimensional images

in any plane (see Figure 1.15). These images are similar to those often acquired post-

operatively, but can be obtained intra-operatively after insertion of the pedicle screws.

The accuracy of the screw placement in the relevant planes can then be assessed. Of

course, three-dimensional fluoroscopy exposes the patient to even greater doses of

ionizing radiation than standard fluoroscopy, since several projection images are required

to create each reconstruction. Ito et al. (2008) described a success rate of 97.2% for

cervical screw insertion using this technique; none of the cortical bone perforations

(2.8%) were clinically significant.

Figure 1.15- Example of intra-operative 3D fluoroscopy imaging. This sagittal reconstruction shows a vertebral body screw placed in the cervical spine. Reconstructions can provide images in axial or coronal planes as well, allowing the placement of the instrumentation to be localized in three-dimensional space. Reproduced with permission from Ito et al. (2008).

21

An alternative to fluoroscopy are infrared-based surgical guidance systems. These

systems generally rely on pre-operative CT or MRI scans for their anatomical

information. In order to provide navigational assistance, these systems utilize infrared

LEDs placed on the vertebrae as well as on the surgical instruments. Infrared cameras

placed around the operating room triangulate the relative position of the surgical

instruments to the anatomy. Figure 1.16 shows images obtained with the Surgical

Navigation Technologies Stealth Station system. Since the anatomical images are

obtained with CT or MRI, reconstructions can be created in any plane, allowing the

pedicle screw to be precisely implanted. Girardi et al. (1999) reported only 3 screw

perforations out of 330 screws implanted in the lumbar spine using this technique; all

perforations were of the lateral cortical wall in narrow pedicles. Bledsoe et al. (2009)

achieved a success rate of 93.3% in upper thoracic spine surgery using a similar

technique; the cortical wall perforations that were observed were low-grade (less than

2mm) and did not lead to clinical symptoms or complications.

Figure 1.16- Example of 3D infrared-based surgical navigation. This 3D reconstruction shows a posterior view of the spine. Important anatomical landmarks (numbered circles)

22

are “registered” with the navigation computer. These points may be registered by attaching an infrared LED or by touching them with an LED-equipped probe. Registration is required in order to properly display the position of the surgical instruments relative to the anatomy. Reproduced with permission and copyright © of the British Editorial Society of Bone and Joint Surgery [Girardi et al. (1999)].

One significant drawback to systems that depend on pre-operative images

becomes apparent if the anatomy changes during surgery. The guidance system may still

verify the best pedicle screw placement based on the pre-operative images; however, the

changes to the anatomy of the spine during surgery are not reflected in these images.

Actual correctness of the pedicle screw fixation based on post-operative CT or MRI scans

may show significant cortical perforation or breach.

1.4 Developing an ultrasound alternative

Ultrasound imaging, as discussed earlier, has been utilized in many fields of

surgery. For spinal fusion guidance, a properly designed ultrasound probe may provide

real-time images, in the relevant directions (superior, inferior, medial, and lateral), and

without ionizing radiation. In contrast to ultrasound probes designed for abdominal or

vascular surgery, an ultrasound probe designed for spinal fusion surgery would be

required to image bone accurately. This requires an understanding of ultrasound

propagation through both trabecular and cortical bone.

1.4.1 Acoustic properties of trabecular and cortical bone

The acoustic properties of bone are significantly different from those of soft

tissues (adipose tissue, muscle, liver, kidney, etc.). There are three major differences:

speed of sound, attenuation, and dispersion.

23

Speed of Sound

The speed of sound of cortical bone is generally in the range of 2500-4000 m/s.

This is approximately twice the speed of sound in most soft tissues, which is generally in

the range of 1450-1645 m/s (Cobbold 2007). The speed of sound in bone marrow is also

close to that of water and soft tissue.

Trabecular bone is a two-phase medium; it is composed of bone and bone

marrow. The bone volume is defined as the percentage of the total volume of the

trabecular bone which is occupied by bone; the remaining volume is occupied mostly by

bone marrow. The bone component of trabecular bone consists of trabeculae, which are

predominantly rod and plate shaped and are interconnected. The space in between the

trabeculae is occupied by bone marrow. In trabecular bone, the speed of sound depends

heavily on the bone volume. As shown in Figure 1.17, the correlation between the bone

volume and the speed of sound is very strong. This relationship has been verified in

simulations as well as experimentally.

In a heterogeneous material such as trabecular bone, where two materials with

different speed of sound exist, fast and slow ultrasound waves may exist and propagate at

different speeds through the trabecular bone. The speed of these individual components,

the fast and slow waves, varies depending on the bone volume, contributing to the overall

effect of speed of sound dependency on bone volume (see Figure 1.18). The amplitudes

of these component waves is also affected by the bone volume, with the amplitude of the

fast wave increasing at higher bone volume while the amplitude of the slow wave

decreases (Mano et al. 2007).

24

Figure 1.17- Correlation between bone volume and S.O.S. in trabecular bone. A greater bone volume (BV/TV) indicates that a larger percentage of the trabecular bone’s volume is occupied by bone. Due to the difference in speed of sound for bone and soft tissue, this increase in bone volume results in an increase in the speed of sound. This trend is evident over a wide range of physiologically realistic bone volume values. Reproduced with permission from Padilla et al. (2005).

Figure 1.18- Fast and slow waves in trabecular bone. The porosity of the trabecular bone is inversely related to the bone volume; a higher porosity indicates a lower bone volume. The effect of the bone volume, or porosity, on the fast and slow wave components in trabecular bone depends largely on the orientation of the trabecular bone to the direction

25

of travel of the ultrasound beam. Many bones in the human body exhibit trabeculae that are mostly parallel to each other. The top graph shows the case where the ultrasound beam’s direction of travel is parallel to the majority of trabeculae; the bottom graph shows the case where the beam is perpendicular to the trabeculae. As the bone volume is increased, the speed of the fast wave increases and the speed of the slow wave decreases; this results in the overall increase of the speed of sound that is observed. Reproduced with permission from Anderson et al. (2008).

Attenuation

The attenuation of ultrasound in bone also differs significantly from attenuation in

other biological tissues. In many soft tissues, attenuation is relatively low, in the range of

0.1-1.57 dB/cm at 1MHz (Cobbold 2007). In cortical bone, attenuation is higher than in

soft tissues, at approximately 2dB/cm at 1MHz. In trabecular bone, attenuation is

considerably higher than in either cortical bone or soft tissues. Specifically, the scattering

component of attenuation is higher in trabecular bone, and dominates the absorption

component. The trabeculae in trabecular bone create complex scattering of the ultrasound

wave at each interface.

The attenuation in trabecular bone is also frequency-dependent. As shown in

Figure 1.19, in both simulations as well as experimental measurements, the attenuation of

ultrasound in trabecular bone increases with increasing frequency. Due to the high

attenuation, most applications of ultrasound in trabecular bone use lower frequencies (in

the range of 1 – 5MHz) compared to the 5 – 20MHz range typically used for soft tissue

imaging. Nonetheless, ultrasound imaging in trabecular bone is still limited to a few

centimetres due to the low penetration of the beam.

26

Figure 1.19- Frequency-dependent attenuation in trabecular bone. As in soft tissues, an increase in the ultrasound frequency results in an increase in attenuation. However, the magnitude of the attenuation in trabecular bone is much higher than in soft tissues. Furthermore, in trabecular bone there is the added variable of the bone volume. For example, at 1MHz, the above graphs show an attenuation of approximately 10 dB/cm (left graph) and 22 dB/cm (right graph). A higher bone volume (right graph) generally results in a higher attenuation; this is due mostly to increased scattering from the trabeculae. Reproduced with permission from Bossy et al. (2007). In trabecular bone, higher bone volume, at least in the typical physiological range,

results in higher attenuation. High bone volume, it is anticipated, will have an effect on

image quality and accuracy.

Dispersion

Most biological tissues exhibit positive dispersion, an increase in the speed of

sound when the ultrasound frequency is increased. Trabecular bone, however,

demonstrates negative dispersion in some cases. The anomalous negative dispersion may

be caused by interference between the fast and slow waves in trabecular bone. Although

each wave component may exhibit positive dispersion in accordance to the Kramers-

Kronig equation, the existence of both wave modes and the resultant interference causes

27

negative dispersion (see Figure 1.20). Bone volume, by affecting the fast and slow waves,

may in turn be a factor in causing either positive or negative dispersion to be observed.

Figure 1.20- Dispersion in trabecular bone. The effect of ultrasound frequency on the speed of sound depends on the speed of the different components, the fast and slow waves. Therefore the type of dispersion observed depends on the porosity or bone volume. From left to right, the above graphs show the dispersion for a vfast (speed of the fast wave) of 1550 m/s, 1600 m/s, 1700 m/s and 2100 m/s. Positive dispersion is shown for the case where the speed of the fast wave is 1550 m/s, while negative dispersion is observed for the two intermediate cases. The final case, in which the speed of sound is highest, exhibits a combination of positive and negative dispersion in different intervals of the frequency range. Reproduced with permission from Anderson et al. (2008).

1.4.2 Design of prototype transducer

Due to the specific acoustic properties of trabecular bone and the shape and size

of pedicles, an ultrasound probe had to be specifically designed for this application. The

prototype probe, shown in Figure 1.21, operates at a centre frequency of 3.2MHz, and

with a diameter of 3.5mm can fit within a cannulated bore hole in a typical thoracic or

lumbar pedicle (Mujagic 2007). Since the probe consists of only a single transducer

element, only one-dimensional (A-mode) images can be obtained; to create two-

dimensional (B-mode) images several consecutive A-mode images would need to be

obtained, with each line in the B-mode image being composed of a single “A-line”. The

direction of movement of the probe between each A-mode image acquisition can be

either along the long axis, or rotation, as shown in Figure 1.22. The first type of

movement, along the long axis, would provide a B-mode image of one side of the

28

pedicle. For example, if the transducer is pointed medially or laterally, an axial image of

the medial or lateral section of the pedicle would be acquired. If the transducer is pointed

cranially or caudally, a sagittal image of the superior or inferior section of the pedicle

would be obtained. Rotation of the transducer allows coronal images of the pedicles to be

created.

Figure 1.21- Prototype ultrasound probe. The transducer element is at the tip of the probe and is approximately 2mm in width.

Figure 1.22- Use of the prototype probe. The experimental setup allows the probe to be either rotated along its axis or translated. Rotation provides coronal images of a cross-section of the object, while translation can provide images in the axial, sagittal or oblique planes depending on the angle the transducer is pointed.

In testing the efficacy of the prototype transducer, the experimental ultrasound

images will be assessed based on the visibility of the cortical bone. This feature is

important if ultrasound imaging is to be used for surgical guidance; images acquired

intra-operatively during the cannulation process should show whether the cortical bone is

intact or if it has been breached, possibly causing neurological or vascular damage.

29

Results of experimental testing of the prototype probe on human cadaveric vertebrae are

discussed in detail in Chapter 3.

1.5 Hypothesis

Our research hypothesis is that, at relatively low frequencies in the range of 1-

3.2MHz, the cortical bone layer should be distinguishable in the ultrasound image,

through several millimetres of water and trabecular bone, over a wide range of bone

volume.

1.6 Research objectives

Our overall objective is to develop an ultrasound imaging system for intra-

pedicular imaging that can be used for guidance of pedicle screw insertion

Our research objectives are:

1) To determine whether, in fact, the cortical bone layer is consistently visible

in ultrasound images.

2) To determine the accuracy of the ultrasound images, specifically regarding

the location of the cortical bone.

3) To elucidate the effect of ultrasound frequency on image quality and

accuracy.

4) To ascertain the effect of bone volume on image quality and accuracy.

For the purposes of the intended application of spinal fusion image guidance,

image quality was defined as the visibility of the cortical bone layer.

30

Chapter 2 – Simulation Methods and Results

Prior to testing the experimental prototype with human pedicles, we sought to

determine the image quality that can be gained using low-frequency ultrasound within

trabecular bone. To achieve this, computer simulations were conducted. The simulation

program, Wave3000, uses a Finite-Difference Time-Domain algorithm to solve the 3D

visco-elastic wave equation, namely

( )wtt

wtt

w•∇∇

∂∂

+∂∂

+++∇

∂∂

+=∂∂

32

2

2 ηφµληµρ (2.1)

The variables in the above equation are material-specific:

ρ – Material density (in kg/m3)

λ – First Lame constant

µ – Second Lame constant

η – Shear viscosity

φ – Bulk viscosity

w – Displacement vector

The simulation program divides the simulation space into a grid, where each

element of the grid is a cube of the same size. The size of the grid elements depends on

the material properties and the simulation frequency. The three-dimensional object to be

imaged (such as bone) is represented by pixels (“voxels”) each of which is also a cube of

the same size. The size of the voxels depends on the resolution of the image of the object,

from which the three-dimensional model is being created. The size of the voxels may be

larger or smaller than the size of the grid elements. Furthermore, the finite simulation

31

time is divided into time-steps. At each time-step, the displacement vector is calculated

using Equation 2.1 for each grid element in the simulation space.

Before conducting simulations, however, the simulation model must be designed.

The simulation model is composed of the geometry, material properties, boundary

conditions, transducers, and simulation parameters.

2.1 Creating a simulation model

In memory- and processor- intensive computer simulations, it is often unfeasible

to model the entire object of interest. Instead, a portion of the object is modelled and

simulated, and the results provide insight into the larger problem. For example, instead of

modelling an entire aircraft wing in a simulation to study the effect of ultrasound on the

wing’s durability, a small portion of the wing may be modelled in simulation, while still

providing results that are applicable to the larger problem. Therefore, the simulation

model must be chosen to be representative of the larger object that is being studied.

The complexity of the simulation problem increases as the size of the object

increases, requiring more memory and more computing time. An increase in the

frequency of the transducer(s) also increases the demands on computer hardware. For

example, doubling the transducer frequency from 1MHz to 2MHz requires a factor of 8

increase in computer memory, since the resolving wavelength is halved in all three

dimensions. These constraints were taken into account in designing the simulation model.

32

Geometry

The geometry consists of the shape of the object(s) to be imaged or sonicated. In

order to create a simulation model which would provide results comparable to

experiments, the anatomy of the thoracic and lumbar pedicles was chosen as the starting

point. As shown in Figure 2.1, the major anatomy in the region being imaged is the

trabecular and cortical bone of the pedicle. Since the diameter of the ultrasound probe is

slightly smaller than the diameter of the bore hole, water is used as a coupling medium

between the transducer and the trabecular bone.

Figure 2.1- Intra-pedicular imaging using ultrasound. After cannulation of the bore hole through the posterior arch and into the pedicle, the ultrasound probe would be inserted into the bore hole and used to image the trabecular and cortical bone (as shown on the left figure). The transducer at the tip of the probe faces a coupling medium composed of water (which is used both experimentally as well as during surgery) followed by a thick layer of trabecular bone and a thin layer of cortical bone (as shown on the right cross-sectional figure). One “slice” of the pedicle contains each of these materials. Left figure modified from İnceoğlu et al. (2007), right figure modified from Moore and Dalley (2006).

For the purposes of simulation, the trabecular bone was perfused with water

instead of bone marrow, which has similar acoustic properties. Since the vertebrae used

33

in experiments are cleaned of all soft tissue, this tissue was also omitted from the

simulation geometry.

The dimensions of each layer of the slice were chosen to be representative of a

typical thoracic or lumbar vertebra. A trabecular bone thickness of 4.9mm was chosen,

along with a cortical bone thickness of 1.0mm. İnceoğlu et al. (2007) reported that the

lateral wall of the pedicle is 609 ± 247 µm and the medial wall 793 ± 186 µm. These

measurements reflect the thickness of the cortical wall at its midpoint, which is the

thinnest section of the cortical wall. The transducer was placed 7.0 mm from the cortical

bone. The transducer was selected to be a 3mm×3mm square, roughly equal in size to the

transducer used in experiments.

Trabecular bone has very complex geometry, consisting of interconnecting

trabeculae, varying in size and alignment. Many attempts have been made at modelling

trabecular bone using, for example, rods of a certain length, width, and orientation to

represent the trabeculae. The exact dimensions and orientation of the rods can be random

(within a specific range) to model the variation in real trabecular bone (Hosokawa 2006).

Alternatively, a parallel-plate model of alternating polystyrene and water layers has been

described by Wear (2001). These models provide similar acoustic characteristics as

trabecular bone while simplifying the task of creating the three-dimensional geometry.

For the purposes of this research, however, the geometry of the trabecular bone was

obtained from a human vertebral specimen scanned by a micro-Computed Tomography

(µCT) scanner. This type of scanner yields several hundred two-dimensional images in

the DICOM file format, which together can be used to reconstruct the three-dimensional

geometry of the trabecular bone. An L4 human vertebral body from a cadaver (F, age 60)

34

was cleaned of all soft tissue and bone marrow, and scanned with µCT, to yield three

segments of 3D trabecular bone geometry. Due to logistical reasons, the first segment

was obtained with a SkyScan 1172 µCT, while the second and third segments were

obtained with a SCANCO Medical 40 µCT. Each segment was selected from a different

anatomical region of the vertebral body.

Since the study of the effect of bone volume on image quality is one of the

objectives of the current research, different bone volume models needed to be created out

of the raw µCT data. The method of trabecular dilation/erosion, shown in Figure 2.2, was

used to create three different bone volume models (low, medium, and high) from each

trabecular bone segment. The range of bone volume was selected to represent the normal

physiological range, from approximately 5% to 15%. Each bone volume model was then

segmented into three slices along its long axis, yielding 27 slices, each with a thickness of

4.9mm. The bone volume of each slice can vary considerably from the model it is derived

from, as shown in Table 2.1. This is due to regional variation in bone volume from one

part of the vertebra to another. The final simulation geometry is shown in Figure 2.3.

Figure 2.2- Dilation and erosion of trabecular bone. The original µCT images are greyscale; a wide range of bone volume can be simulated through different assignments of voxels (by colour) to different materials. Dilated trabeculae (left image) exhibit larger trabecular thickness and lower spacing between trabeculae than eroded trabeculae (right image). The general geometry of the trabecular bone, such as the orientation of the trabecular rods, is preserved.

35

TABLE 2.1- BONE VOLUMES FOR MODELS AND INDIVIDUAL SLICES

Segment Size, Resolution

Bone volume, each

model

Bone volume, individual slices

Slice #

SkyScan 1172 3.3×18.3×3.5mm

57.3 voxels/mm or 0.017 mm/voxel

6.7% (low)

5.4% 7.7% (low)

9.0%

1 2 3

11.4% (mid)

9.5% 12.9% (mid)

15.0%

4 5 6

15.0% (high)

12.8% 16.9 % (high)

19.2%

7 8 9

SCANCO Medical 40 3.0×17.0×3.0mm

66.7 voxels/mm or 0.015 mm/voxel

6.9% (low)

5.2% 10.5% (low)

6.0%

10 11 12

11.4% (mid)

9.0% 15.7% (mid)

11.1%

13 14 15

15.5% (high)

11.9% 21.5% (high)

14.8%

16 17 18

SCANCO Medical 40 3.0×15.0×3.0mm

66.7 voxels/mm or 0.015 mm/voxel

6.8% (low)

4.7% 8.3% (low)

7.1%

19 20 21

11.7% (mid)

9.2% 13.2% (mid)

12.3%

22 23 24

15.5% (high)

12.4% 17.3% (high)

16.4%

25 26 27

36

Figure 2.3- General simulation geometry. The trabecular bone is oriented such that the majority of the parallel trabecular rods are parallel to the direction of propagation of the plane wave, allowing the worst-case attenuation to be simulated. Each slice of trabecular bone is 4.9mm in thickness. The cortical bone thickness was selected to be 1.0mm; the cortical bone is oriented perpendicular to the direction of propagation of the ultrasound wave. The total distance between the transducer and the proximal wall of the cortical bone was 7.0mm.

Material Properties

In order to create a simulation model that is comparable to the experimental

protocol, soft tissue (such as bone marrow, muscles, and adipose tissue) was not

included. Therefore the simulation model contains two distinct materials: water and bone.

The material and acoustic properties of water (density, speed of sound, absorption,

viscosity, etc.) are well documented. For cortical bone, many studies have been

conducted to elucidate these properties. The values provided by Luo et al. (1999) and

Wear (2001) are identical to those in Wave3000 for a 1MHz transducer. Prior studies

37

(Mujagic 2007) have modelled trabecular bone as a two-phase medium: the trabeculae

are assigned the material properties of cortical bone, and the space between the

trabeculae, normally occupied by bone marrow, is replaced by water. These models have

yielded values for the speed of sound and attenuation that accurately reflect measured

values in trabecular bone specimens. Table 2.2 shows the material properties of bone and

water used in simulation, including the wave speed and wavelength. By comparison,

trabeculae are generally 0.043-0.110 mm in width and 0.50-1.34 mm apart (Chaffaî et al.

2002). A typical slice would be assigned material properties as shown in Figure 2.4.

TABLE 2.2- MATERIAL PROPERTIES Parameter Bone Water

Density 1850 kg/m3 1000 kg/m3

Longitudinal wave speed

2901 m/s @ 1MHz 2902 m/s @ 2MHz

1497 m/s @ 1MHz 1497 m/s @ 2MHz

Transverse wave speed

1303 m/s @ 1MHz 1313 m/s @ 2MHz

-

Longitudinal wavelength

2.901 mm @ 1MHz 1.451 mm @ 2MHz

1.497 mm @ 1MHz 0.748 mm @ 2MHz

Transverse wavelength

1.303 mm @ 1MHz 0.656 mm @ 2MHz

-

Shear viscosity 40 Pa·s - Bulk viscosity 0.1 Pa·s 1 mPa·s

First Lame constant

9306 MPa 2241 MPa

Second Lame constant

3127 MPa -

Boundary Conditions

The 3D simulation space has six boundary surfaces. Early simulation work on this

project utilized infinite boundary conditions on all sides of the simulation space;

however, the resulting unmitigated diffraction effects often led to unstable simulations.

Use of limiting boundary conditions allows a plane wave to be transmitted. Boundary

38

conditions that restrict wave motion were used for the four sides whose normal vector is

perpendicular to the direction of the transmitted plane wave. The remaining two sides

were selected to have infinite boundary conditions.

The infinite boundary condition furthest from the transducer serves to prevent

reflections from beyond the cortical bone. Therefore, the echo signal can be guaranteed to

originate from either the trabeculae or the cortical bone. The portion of the ultrasound

wave which does not reflect from the bone interfaces is absorbed by the infinite boundary

condition.

Figure 2.4- Example of material properties assignment. Voxels assigned the material properties of bone are shown in red, and voxels assigned the properties of water are shown in green. Due to the absence of bone marrow, the acoustic impedance of the trabecular bone would be expected to be closer to that of water. The trabeculae and the cortical bone are assigned identical material properties; therefore, the differences in attenuation and speed of sound between the two tissues will be determined primarily by scattering (i.e., by the trabecular bone geometry) rather than by absorption.

39

Transducers

Higher transducer frequencies, with their shorter wavelengths, require a smaller

spatial resolution and therefore increase the computation time. Due to this factor, the

1MHz simulations were conducted first, followed by the 2MHz simulations. The

transmitted wave was a single wideband sinusoidal pulse, with amplitude of one. The

transducer, a 3mm×3mm square, was given a void backing to prevent backward

transmission or reflection artifacts. The receiving function of the transducer was delayed

by 1µs from the start of the simulation to prevent overlap between transmission and

reception.

Since the transducer occupies an entire surface of the simulation space, diffusion

effects at the edges of the transducer are nonexistent. The transmitted pulse has the shape

of a plane wave, as shown in Figure 2.5.

Simulation Parameters

Several simulation parameters in Wave3000 allow for fine-tuning the simulation.

The major factors affected by these parameters are the simulation speed, accuracy, and

stability.

The simulation time is the total duration of the simulation (in microseconds). This

includes the time required to transmit the pulse as well as the time to receive echoes from

the farthest objects in the simulation space. Given that in our simulation model, the

cortical bone (at 7.0mm) is the most distant reflector from the transducer, an estimate of

the speed of sound in the trabecular bone can provide a preliminary figure for the

40

required simulation time. Increasing the simulation time results in an increase in the time

required to complete the simulation.

Figure 2.5- Transmitted plane wave. The positive and negative peaks of the single transmitted pulse are shown as white bands. Since no apodization factor was assigned for the transducer, the magnitude of the pressure is constant along the wave. Reflections from the trabecular or cortical bone, however, are not necessarily plane waves.

The time-step directly affects the speed and stability of the simulation. A higher

time-step allows for faster simulations by reducing the number of computations during

the simulation run. For example, for a 10µs simulation consisting of 1000 time-steps

(0.010µs per step), a decrease of the time-step to 0.0060µs would increase the total

number of steps in the simulation run to 1667. This may be required if the simulation

41

proves unstable at the higher time-step. However, reducing the time-step increases the

computation time.

Spatial parameters can also be selected to increase or decrease the accuracy of the

simulation. These parameters depend largely on the frequency of the transducer and the

materials properties in the simulation model. In our simulation model, the material in

which the shortest wavelength is present is water. At 1MHz, the wavelength (assuming a

speed of sound of 1497 m/s) is 1.497mm; at 2MHz the wavelength is reduced to

0.748mm. The resolving wavelength is selected to be shorter than these wavelengths. In

addition to the resolving wavelength, the points-per-cycle parameter determines the size

of the simulation grid. A higher points-per-cycle results in a larger number of grid

elements per wavelength and more accurate (and stable) simulations.

Table 2.3 summarizes the selected simulation parameters for the 1MHz and

2MHz simulations.

TABLE 2.3- SIMULATION PARAMETERS

Parameter Value Time-step 0.0010 µs

Resolving wavelength 0.50 mm @ 1MHz 0.25 mm @ 2MHz

Simulation time 16 µs Points-per-cycle 20

Grid size 0.0250 mm @ 1MHz 0.0125 mm @ 2MHz

Grid/voxel ratio For SCANCO µCT: 0.60 @ 1MHz 1.20 @ 2MHz

For SkyScan µCT: 0.70 @ 1MHz 1.40 @ 2MHz

42

2.2 Simulation analysis and results

Each of the 54 simulations yields a single A-line image, such as the ones shown

in Figure 2.6. The blue curve shows the transmitted ultrasound pulse, and the red curve

shows the received echoes from the trabecular and cortical bone. Looking at this time-

domain data for several slices, two major features become noticeable. The earlier echo

varies in magnitude from slice to slice, but generally does not experience any time-shift,

i.e. it appears at the same time in the simulation. The later echo also varies considerably

in magnitude, but also displays a visible time-shift. By inspection, it also appears that the

frequency of the second echo is significantly lower than that of the transmitted signal or

the first echo.

The most pertinent question is the accuracy of the distance measurement from the

transducer to the cortical bone, and the dependency of the accuracy on the bone volume

or transducer frequency. In order to answer this question, the time-domain simulation

results must be converted into A-mode distance graphs. Considering that the speed of

sound in bone varies considerably depending on the bone volume, constructing truly

accurate images would require a priori knowledge of the bone volume. However,

ultrasound images are generally constructed by assuming a fixed speed of sound,

typically 1545 m/s. This assumption was used in our image analysis in order to determine

whether a fixed speed of sound would yield sufficient image accuracy over a wide range

of physiological bone volume, or whether the speed of sound would need to be selected

depending on an estimate or exact measurement of the bone volume.

43

Figure 2.6- A-line simulation ultrasound images. The earliest echo (A) originates from the water/trabecular bone interface at 2.1mm, while the later echo (B) originates from the trabecular/cortical bone interface at 7.0mm. Since the two components of the received signal have propagated through different media, they exhibit different frequency spectrums.

44

2.2.1 Image accuracy

For each simulation, the time-of-flight was computed as the time between the first

peak of the cortical bone reflection (β) and the first peak of the transmitted ultrasound

wave (α).

Using the assumed speed of sound of 1545 m/s, the distance can be calculated as

Measured Distance = 2

flightofTime −−× Assumed Speed of Sound (2.2)

This is the measured distance between the transducer and the cortical bone. Next, the

error (in millimetres) between the measured and actual distance, as well as the

dependency of this error on bone volume and transducer frequency, was investigated.

Figure 2.7 shows the error vs. bone volume for both 1 and 2MHz simulations.

Higher bone volume generally results in a higher speed of sound (due to the

preponderance of fast waves), which reduces the time-of-flight. The shorter time-of-flight

results in a smaller measured distance. This can be seen clearly in Figure 2.8, which

shows the measured distance vs. bone volume for all slices. A selection of a different

speed of sound would cause a shift in the y-axis of Figure 2.8. The assumed speed of

sound underestimated the distance of the cortical bone for most slices; a higher assumed

speed would yield better average accuracy for a wide range of bone volume.

45

Figure 2.7- Error vs. bone volume and frequency. The error between the measured and actual distance depends on the assumed speed of sound. A higher assumed speed of sound can reduce the error for high bone volume slices while increasing the error for low bone volume slices, therefore reversing the observed trend.

Figure 2.8- Measured distance vs. bone volume and frequency. Each data point represents a particular slice simulated at one frequency. The existence of fast and slow waves therefore directly affects the accuracy of the ultrasound measurements.

46

Neither Figure 2.7 nor Figure 2.8 shows a significant dependency of the measured

distance or distance error on transducer frequency. In some slices, an increase in the

transducer frequency results in an increase in measured distance, i.e., an increase in the

time-of-flight. This implies that the speed of sound has decreased (negative dispersion).

This is true for many of the low-bone volume slices. Amongst the high-bone volume

slices, some cases of positive dispersion can be seen. However, the dependence of the

distance measurement accuracy on transducer frequency is small compared to the

dependence on bone volume, which is the primary factor affecting the distance

measurement accuracy.

In order to study the qualitative effects of transducer frequency and bone volume

on image quality, colour A-mode images were created for each A-line graph. These

images were created by using a Hilbert transform to calculate the envelopes of the echo

signals from the A-line graphs. Figure 2.9 shows a collection of 14 colour images from

different slices. The horizontal axis shows the distance from the transducer, which is

located at the origin. The colour axis shows the intensity of the echoes. In most images,

the water/trabecular bone interface at 2.1mm and cortical bone at 7.0mm can be

distinguished relative to the background reflections caused by individual trabeculae. The

relative intensity of the echoes from those important interfaces, relative to each other as

well as relative to the smaller background echoes, depends to a great extent on the bone

volume and the transducer frequency.

47

Figure 2.9- Simulation A-mode images. The transducer is placed at the origin. While the 2MHz images are sharper, the cortical bone is less visible. In some slices, such as slice #8, the cortical bone is barely identifiable at high frequency relative to the background echoes, while the trabecular bone has become more visible. In some slices, such as slice #4, the trabecular bone interface is not visible; this is due to the particular trabecular geometry of this slice, in which trabecular spacing or alignment is different than other slices of the same bone volume. The reflections from the cortical bone also appear “wider” than the reflections from the trabecular bone; this is due to the frequency-dependent attenuation which reduces the centre frequency of the cortical bone echo.

48

Low bone volume means that the water/trabecular bone interface is more porous

and has an acoustic impedance close to that of water; this allows the ultrasound beam to

easily penetrate this interface while the bone causes only a weak reflection. If the

transducer frequency is increased, however, the trabecular bone interface generally

becomes clearer, as the higher-frequency ultrasound beam is scattered more by the

smaller trabeculae. However, this results in less ultrasound energy penetrating through

the trabecular bone interface. This factor contributes to the reduced visibility of the

cortical bone at high frequency. The other major factor contributing to the reduced

visibility of the cortical bone is the scattering and absorption caused by the trabecular

bone itself. Therefore, despite the improved sharpness of the 2MHz images, compared to

the 1MHz images, the increased scattering and absorption result in lower visibility of the

cortical bone.

As the bone volume is increased, the relative visibility of the trabecular and

cortical bone interfaces is interchanged. The cortical bone becomes less visible due to the

higher attenuation (mostly scattering) from the trabecular bone, while the trabecular bone

interface becomes more visible.

Furthermore, due to the effect of fast waves, the cortical bone is shifted in the

high bone volume images. The cortical bone therefore appears closer to the transducer

than it does in images of low bone volume slices. For some slices, such as slice #26 in

Figure 2.9, negative dispersion partly offsets the increased speed of sound that is due to

the higher bone volume. The result is that, for some slices, the distance shift is not as

pronounced at 2MHz as at 1MHz.

49

2.2.2 Attenuation

As shown in Figure 2.9, high bone volume causes a qualitative diminishing of the

visibility of the cortical bone, especially at high frequency. In order to quantitatively

examine the effect of bone volume and transducer frequency on attenuation, it is required

to isolate the signal of the cortical bone from the original A-line graph. Where possible,

the closest x-axis interception before the cortical bone reflection was used as the

beginning of the sampling window. The sampling window ends at the end of the

simulation time. This serves to isolate the cortical bone signal from the trabecular bone

echoes, which having travelled a shorter distance in the trabecular bone, have a different

frequency spectrum than the echo from the cortical bone.

The frequency spectrum of the cortical bone echo was obtained using the Fast

Fourier Transform. In order to obtain the exact centre frequency of the transmitted pulse,

its FFT was calculated; the actual values used in later calculations were 1.1MHz and

1.8MHz for the 1MHz and 2MHz calculations, respectively. The attenuation for each

slice at each frequency was calculated from

))(

)(log(20)(

fA

fAf

ref

sig=α (2.3)

The results were grouped according to bone volume and transducer frequency. Averages

and standard deviations were calculated for each bone volume category (low, medium,

and high). The results shown in Figure 2.10 reaffirm the qualitative trends shown in the

colour A-mode images: attenuation is higher for high bone volume slices and increases as

the transducer frequency increases. The standard deviations provide insight into the effect

that variations in trabecular bone geometry and bone volume (within the same bone

50

volume category) have on attenuation. These variations between slices of the same

category can be considerable, but are less significant than the difference between separate

categories. The major factor causing these variations is the particular trabecular geometry

of the bone sample from which the bone volume models and slices are created. In

selecting the bone segments from the original µCT scans, care was taken to avoid

segments with visible anomalies such as especially porous areas (which can be common

in bone samples from older donors). A bone sample from a different donor, with a

different bone volume, may have more variation in regional bone volume or trabecular

geometry, which would be reflected in A-mode colour images and attenuation figures.

Figure 2.10- Attenuation dependence on bone volume and transducer frequency. The error bars show one standard deviation. N=9 for each category.

Of the two component parts of attenuation (scattering and absorption), scattering

is the dominant contributor at these frequencies. The water/trabecular bone interface

contributes profoundly to the high attenuation of the signal that reaches the cortical bone.

The simulation results suggest that a lower transducer frequency would be

preferred for imaging due to the lower attenuation. The improved resolution of the high

51

frequency images would not be useful in surgical guidance if the cortical bone is not

visible.

The accuracy of the simulation images obtained suggests that the original

assumption of a constant speed of sound is a valid assumption over a wide range of bone

volume. Although the assumed speed of sound can be improved to yield better average

accuracy over the entire physiological range of bone volume, a fixed speed of sound can

still be used, eliminating the need for prior knowledge of the bone volume of the

vertebra.

52

Chapter 3 – Laboratory Experiments

The simulation results provided insight into the effect of bone volume and

transducer frequency on image quality. They also affirmed that the cortical bone should

be visible even in the higher end of the physiological bone volume range, at transducer

frequencies up to 2MHz and through as much as half a centimetre of trabecular bone.

Laboratory experiments introduce new variables, such as trabecular and cortical

bone thickness and cortical bone curvature. These variables depend on the particular

morphology of the vertebral bone specimen and the spine level (cervical, thoracic, or

lumbar). There is also variation within the same bone specimen; for example, the medial

cortical wall of the pedicle is thicker than the lateral, and the diameter of the pedicle (and

therefore the thickness of the trabecular bone) varies along the length of the pedicle.

Larger vertebrae have larger pedicles and therefore greater trabecular bone

thickness. The increased trabecular bone thickness causes higher attenuation and

therefore may reduce the visibility of the cortical bone. In order to investigate the worst-

case scenario, large lumbar vertebrae were used in the laboratory experiments.

Unlike computer simulations, experiments must be conducted with a constant

bone volume; the bone volume depends on the age of the donor as well as the incidence

of osteoporosis or other ailments that can affect bone volume. The bone specimens used

in these experiments were obtained from older donors, which may have bone volume in

the lower end of the physiological range or may be partially osteoporotic. Low bone

volume, as shown by the computer simulations, should improve visibility of the cortical

bone.

53

The optimum transducer excitation frequency depends on the thickness of the

piezoelectric transducer medium. As described in section 1.4.2, the centre frequency of

the prototype transducer used in these experiments is 3.2MHz. At this frequency, it was

anticipated that the water/trabecular bone interface would reflect much of the incident

ultrasound energy, possibly hampering imaging of the cortical bone.

3.1 Experimental protocol

Vertebral bone specimens were acquired from the Anatomy Department at the

University of Toronto. The specimens came with soft tissue, including muscle, adipose

tissue, ligaments, and bone marrow. In order to remove the soft tissue, a scalpel was used

to clean the exterior of the vertebrae, while a water jet was used to clear the bone marrow

from within the vertebrae. Finally, a methanol solution was used to clean the specimens

(Mujagic 2007). Due to the complex morphology of vertebrae, complete removal of the

soft tissue was very difficult; therefore, the bone specimens were stored in a freezer for

preservation. This prevents bacteria from growing on the soft tissue.

Each vertebra was divided into two parts along the medial plane using a bone

saw. The pedicles were then isolated from the vertebral body and the posterior arch.

Isolation of the pedicles facilitated cleaning of the bone marrow and other soft tissue near

the pedicle. Identification of the “entry point” for the cannulation probe was aided by the

removal of most of the posterior arch. The cannulation probe was then utilized to bore a

hole with a diameter of approximately 4mm through the centre of each pedicle.

For imaging, each pedicle was placed in a three-point holder, with the pedicle

aligned with the posterior end at the top and the anterior end at the bottom, as illustrated

54

in Figure 3.1. The pedicle and holder were then immersed into the water tank shown in

the right-hand side of Figure 3.2, and held rigidly to the water tank’s frame to prevent

motion.

Figure 3.1- Specimen placed in three-point needle holder. The posterior face of the pedicle with the 4mm-wide bore hole is shown.

The transducer was then placed into a separate holder connected to four individual

motors that allow for movement of the transducer in four directions: X, Y, and Z axis,

and rotation about the Z axis. The pedicle was then manually aligned to enable the

transducer to be placed inside the bore hole. This resulted in the transducer being

approximately 2mm from the trabecular bone surface, a similar distance to that used in

the simulations. The transducer was connected to a Panametrics Model 5800

transmitter/receiver operating with a 200Hz repetition frequency. The receiver output

signal was sent to a Tektronix TDS3012B oscilloscope, which is connected via an

Ethernet cable to a desktop PC. Custom designed software captures the oscilloscope

signal and stores it, and controls the four stepper motors.

55

Since the transducer consists of a single piezoelectric element, one-dimensional

A-line images can be created, and displayed on the oscilloscope. In order to create two-

dimensional B-mode images, multiple A-line images must be obtained at regular

intervals. In these experiments, coronal images (cross-sectional images of the pedicle)

were created; therefore, the direction of motion of the transducer between each A-line

image acquisition was a rotation about the Z-axis. The minimum rotation angle allowed

by the hardware setup (approximately 5.4 degrees) was used. A reference A-line image,

obtained with the transducer in the tank without the bone specimen, was also acquired

before each experimental run.

Figure 3.2- Experimental equipment, including the oscilloscope, transmitter/receiver, and 130 litre water tank. Distilled water was used to couple the ultrasound transducer to the specimen, while minimizing possible contaminants.

56

Due to the variation in the shape of the pedicle along its length, cross-sectional

images should be acquired at different regions of the pedicle. Each cross-sectional image,

comprising a 360° view of a specific region of the pedicle, consisted of 67 A-line images.

A MATLAB script was used to combine each group of A-line images into a single B-

mode image, utilizing low-pass interpolation for smoothing. The ultrasound images were

compared to µCT images or photographs of the same pedicle. The first specimen was an

L4 lumbar vertebra (M, age 83), while the second specimen was an L1 (M, age 88).

3.2 Analysis of ultrasound images

Figure 3.3 shows two A-line images as they are displayed on the oscilloscope.

Each A-line image contains both the transmitted pulse and the received echo. In order to

isolate the received echo, the reference (water only) signal was subtracted from each A-

line image. A comparison between several A-line images shows that the first 2 µs

(approximately) are identical among all images and the reference signal; this time frame

contains the transmitted pulse. The magnitude and location of the received echoes varies

widely between A-line images. This can be observed during experiments as the

transducer is rotated through its 360° course: distinct echoes appear as the transducer is

rotated and gradually disappear as the transducer is turned away from the scatterer (which

can be cortical bone or trabeculae).

In order to determine the source of the scattering, the time-domain A-line graphs

must be converted into distance A-mode and finally colour B-mode images. It was

assumed that the effective speed of sound was 1545 m/s.

57

Figure 3.3- Experimental A-line ultrasound images. The y-axis shows voltage (in Volts), and the x-axis shows time (in microseconds). The total duration is approximately 10µs. Separation of the transmitted and received signals is accomplished by subtracting the reference signal from the A-line image.

58

Despite the subtraction of the reference signal from each A-line image, the first

several microseconds of any A-line image may still contain faint signals which would

appear as echoes in a colour image. The centre of each B-mode cross-sectional image,

however, is occupied by the transducer itself and echoes that appear in this region are

caused solely by noise. Likewise, a small area around the transducer contains water and

should not return any reflections. Since the exact diameter of the transducer and the bore

hole are known, these regions can be blacked-out to improve the clarity of the images.

Finally, each group of 67 A-mode images can be combined to form a single 360°

B-mode image. Two images from the L4 vertebra are shown in Figure 3.4. Outside of the

4mm-diameter centre region, reflections from various scatterers can be seen. The

intensity of the echoes (as shown by the colour bar) is highest in the medial and inferior

directions and dimmer in the lateral and superior directions. Unlike most simulation A-

mode images, which show a distinct reflection from each of the two bone interfaces,

experimental images generally show at most one strong reflection from any given angle.

The close proximity of these visible echoes to the transducer and the high centre

frequency of the transducer suggest that the scattering is more likely caused by the

water/trabecular bone interface than by the cortical bone layer.

59

Figure 3.4- B-mode ultrasound images of left pedicle of L4 vertebra. Top image was acquired with the ultrasound probe placed just inside the pedicle, while the bottom image was acquired after the probe had been advanced 8mm anterior (towards the vertebral body).

To deduce the source of the scattering, µCT images were obtained for the above

pedicle using the SCANCO µCT scanner at a resolution of 55.6 voxels/mm. Since each

µCT slice is 0.018mm thick and the ultrasound transducer has a length of 4mm, each

60

ultrasound image actually encompasses many consecutive µCT slices. For the purposes

of comparison, the sixty µCT slices closest to the centre of the transducer were selected.

A two-dimensional projection of these slices was produced by using the Maximum

Intensity Projection (MIP) feature of the software program GE MicroView. Since a MIP

image illustrates a larger region of the pedicle, it provides a better comparison with an

ultrasound image than a single µCT slice. The corresponding µCT and MIP images for

the ultrasound images in Figure 3.4 are shown in Figure 3.5.

Figure 3.5- Micro-CT (left) and Maximum Intensity Projection (MIP) images (right) of the left pedicle of the L4 vertebra. Top images correspond to the first ultrasound image in the previous figure; bottom images correspond to the ultrasound image obtained 8mm anterior. MIP images highlight the trabecular geometry over a region of the pedicle, while the µCT images show the trabecular geometry near the centre point of the transducer.

61

The µCT images show that the shape of the pedicle varies considerably along the

axis of the pedicle. In the first MIP image shown in Figure 3.5, the cortical bone layer in

the medial and inferior directions is approximately 8mm and 4mm away from the centre

of the bore hole, respectively. The reflections in the first ultrasound image are therefore

more likely caused by the trabecular bone than the cortical bone, even when the tendency

for ultrasound measurements to underestimate the actual distance (as suggested by the

simulation results) is taken into consideration. The second MIP image shows the medial

and inferior cortical walls are still roughly the same distance from the centre of the bore

hole. The corresponding ultrasound image shows a strong reflection from the inferior

direction, at a distance of approximately 4mm from the centre of the transducer, most

likely caused by the inferior cortical bone layer. A possible factor allowing the cortical

bone to be visible in the second image may be higher porosity of the trabecular bone in

this region, which would contribute to improving the visibility of the cortical bone by

reducing attenuation. The trabecular spacing in the second µCT image is noticeably

larger than in the first image. The curvature of the superior and inferior cortical walls is

also different between the two regions of the pedicle. Finally, the second MIP image

shows that the trabeculae are mainly aligned in the cranio-caudal direction in that region

of the pedicle. These trabeculae are parallel to the medial cortical bone layer and are

likely the source of the reflections seen in the medial direction in the second ultrasound

image.

The right pedicle of the same L4 vertebra was also imaged with the prototype

ultrasound probe. The probe was advanced to near the midpoint of the pedicle, between

the posterior arch and the vertebral body. After imaging, the pedicle was bisected along

62

the imaged plane (Figure 3.6). The photographs show that the cannulation of the bore

hole in this case was close to the superior and lateral cortical walls rather than along the

preferred centre of the pedicle. The B-mode ultrasound image likewise shows more

visible scattering from the superior and lateral regions. According to the ultrasound

image, the scatterers are approximately 3mm from the centre of the transducer. In the

other directions where the scattering is less pronounced, the echoes are brightest closest

to the transducer and diminish to the level of background noise within 1mm. Due to the

proximity of the bore hole to the superior and lateral cortical walls, the thickness of the

trabecular bone in those regions is very small.

The L1 lumbar vertebra was imaged in the same manner. In this case, the images

were obtained where the pedicle joins the vertebral body. Since the vertebral body was

removed during cleaning and isolation of the pedicle, if the ultrasound probe is advanced

out of the pedicle the trabecular and cortical bone should cease to be visible. This

scenario simulates an inadvertent breach of the cortical wall during cannulation. Figure

3.7 shows B-mode images acquired for the right pedicle of L1. Within the confines of the

pedicle, the ultrasound probe shows reflections from all directions, especially in the

medial and lateral directions. The signal to noise ratio appears sufficient to differentiate

trabecular and cortical bone from background noise.

63

Figure 3.6- B-mode ultrasound image and photographs of right pedicle of L4 vertebra. After the ultrasound image was obtained, the pedicle was bisected along the coronal plane corresponding to the centre point of the transducer. The left photograph shows the anterior face of the posterior section, and the right photograph shows the posterior face of the anterior section. In this region of the pedicle, the bore hole was closer to the superior and lateral cortical walls. The most visible echoes in the ultrasound image are from the lateral and superior directions.

64

Figure 3.7- B-mode ultrasound images of right pedicle of L1 vertebra. As the ultrasound probe is advanced out of the pedicle, the image of the trabecular and cortical bone starts to diminish. The ultrasound probe was advanced 1mm between each successive image, running from the top left to the bottom right.

65

Overall, the experimental results suggest that when the cortical bone is visible, the

accuracy of the distance measurement is excellent. However, in most cases, significant

scattering from the trabecular bone hinders imaging of the cortical bone. Two possible

causes for high visibility of the trabecular bone and low visibility of the cortical bone are

indicated by the simulation results: high bone volume and high transducer frequency.

Since the vertebral specimens used in the above experiments were obtained from older

donors, the bone volume was most likely not in the higher end of the physiological range.

The high centre frequency of the transducer is therefore the probable cause of the high

scattering.

During surgical guidance, the ultrasound images would be used not just to image

the bone but also to provide guidance for pedicle screw insertion. The first two

ultrasound images would suggest to the surgeon that the medial and inferior cortical walls

are the closest to the transducer, while in fact the medial wall is quite distance and the

superior wall is slightly closer than the inferior wall. The third ultrasound image, showing

the superior and lateral cortical walls being closest to the transducer, would provide the

surgeon with the correct guidance so that the trajectory of the pedicle screw can be

adjusted to be more medial and caudal. Finally, the last ultrasound images would suggest

that the ultrasound probe has breached the boundaries of the pedicle and an alternative

screw trajectory needs to be employed. For the ultrasound images to be used safely for

surgical guidance, further study will need to be conducted to investigate the factors that

influence visibility of the trabecular and cortical bone and to provide surgeons with an

understanding of the challenges of ultrasound propagation in bone necessary to correctly

interpret and utilize these images.

66

Chapter 4 – Summary and Conclusions

Some previous studies have examined the quantitative effects of ultrasound

propagation using the through-transmission technique. These studies provided a valuable

understanding of the effects of absorption, scattering, and dispersion on ultrasound beams

travelling through trabecular bone. In order to study the qualitative effects of these

phenomena on image quality, A-mode and B-mode images were obtained using both

simulation and experimental protocols. The starting point was the anatomy of the spine

and consideration of the specifics of spinal fusion surgery. The effects of bone volume,

transducer frequency, and pedicle morphology on image quality were studied. The results

affirmed the feasibility of using ultrasound to image bone with sufficient accuracy for

surgical guidance.

4.1 Summary of contributions

Simulation programs utilizing Finite-Difference Time-Domain methods have

been used to study the effects of trabecular bone on ultrasound propagation and to

compare these results to those observed through experiments, and the accuracy of this

method has been confirmed by numerous authors (Bossy et al. 2007, Padilla et al. 2005,

Kaufman et al. 2008). The current research employed this method to create colour A-

mode images and to study the effect of bone volume and transducer frequency on these

images. The quantitative and qualitative effects of attenuation, dispersion, and speed of

sound on image accuracy were also determined. In general, the image quality was

sufficient to identify the cortical bone even for high bone volume and high frequency

simulations. The simulation results suggested, however, that as the transducer frequency

67

and bone volume continue to be increased, the cortical bone would eventually be

unidentifiable compared to the background scattering.

Earlier applications of ultrasound measurements in bone consisted of measuring

the speed of sound, and deducing a bone volume from that measurement. Previously

conducted experiments regarding ultrasound imaging of bone used vertebral bodies and

piston transducers with an aluminum plate acting as the cortical bone reflector. The

experiments described above with lumbar pedicles introduce difficulties due to the

morphology of the pedicles; curvature of the cortical bone, variation in the trabecular

bone thickness, and variation in the trabecular bone geometry result in significant

inconsistency between ultrasound images. However, the general quality of the ultrasound

images substantiates further experimental work. The most important question to be

answered by future research is why the trabecular bone is not consistently visible. Ideally,

both the trabecular and cortical bone interfaces would be visible and differentiable from

each other. Future experiments should focus on using cervical vertebrae in order to

determine whether the reduced pedicle diameter and trabecular bone thickness will aid in

imaging of the cortical bone. Image processing techniques such as time-gain

compensation should be utilized to partially compensate for the high attenuation caused

by the trabecular bone. It is also possible that the frequency spectrum of the reflections

can be used to determine their source; reflections from the cortical bone should have a

lower centre frequency due to the attenuation of the higher frequencies by the trabecular

bone.

68

4.2 Limitations

One of the main limitations of the simulation protocol used in this research is the

fact that the bone interfaces are flat, resulting in a sudden change in acoustic impedance

between the water, trabecular bone, and cortical bone. While an abrupt interface between

the water and trabecular bone is realistic, the transition from trabecular to cortical bone is

in reality more gradual. This may result in lower scattering from the cortical bone layer.

Furthermore, the cortical bone was modelled as a homogenous medium, while

anatomically the cortical bone consists of individual osteons, whose size is somewhat

larger than that of the trabeculae. The osteons, which are packed closely together, are

cylindrically shaped and might contribute to scattering.

Another limitation is the fixed trabecular bone thickness employed in the

simulations. Although the effect of trabecular bone thickness on attenuation is

predictable, the effect on distance measurement accuracy has not been examined.

Theoretically, an increase in the trabecular bone thickness would increase the discrepancy

between the actual and measured distance. However, for trabecular bone thickness in the

normal range for cervical, thoracic, and lumbar vertebrae, this effect is relatively

insignificant.

Experiments are more difficult to reproduce than simulations, and introduce

variables such as noise that are not present in simulations. The advantage of experiments

is that they can be conducted relatively quickly and take into consideration the shape of

various pedicles, whereas simulation models must be simplified to adhere to

computational constraints and require significant computation time.

69

One complication of the experimental work arises during comparison between

ultrasound and µCT or photographs. While the µCT images or photographs provide

information about the distance from the centre of the bore hole to the cortical bone, the

ultrasound images provide a measurement of the distance from the transducer to the

cortical bone. The ultrasound probe’s location within the bore hole is not exact. However,

given that the ultrasound probe is only slightly (0.5mm) smaller than the bore hole, the

assumption that the transducer is located approximately in the centre of the bore hole is

valid.

Trabecular bone does not exhibit a constant bone volume throughout a bone;

instead, the bone volume may be higher close to the cortical bone due to the gradual

transition from trabecular to cortical bone. The increase in bone volume near the cortical

bone creates another scatterer that may be confused with the cortical bone but is in fact a

few millimetres closer to the transducer. Additionally, it is possible that pockets of bone

marrow may cause scattering, but due to the similar acoustic properties with water, this

scattering would be minimal. The experimental protocol can be modified to include

complete removal of the posterior arch and exposure of the trabecular bone of the pedicle

to facilitate better cleaning of the bone marrow.

Finally, most specimens are obtained from older donors with low bone volume

and partially osteoporotic bone. In reality, patients with this type of bone would not be

good candidates for spinal fusion surgery; the pedicle screws would not be able to

achieve sufficient grip with the extremely porous trabecular bone. Therefore, most

patients would likely have higher bone volume, which would result in higher attenuation

and reduced image quality.

70

4.3 Conclusions

As stated in the research objectives, we have investigated the effect of bone

volume and transducer frequency on image quality, namely the visibility of the cortical

bone using our simulation protocol. The correlation between high bone volume, meaning

high trabecular thickness and low trabecular spacing, and attenuation and speed of sound

was confirmed. The effect of the bone volume and transducer frequency on the accuracy

of the measured distances was also examined. The image quality anticipated by the

simulation results was superior to that observed experimentally; the cortical bone was

often not visible in experimental B-mode images. The high frequency of the prototype

probe, combined with the large pedicle diameters, results in high attenuation and hampers

imaging of the cortical bone. The complex shape of lumbar pedicles complicates the task

of image analysis and interpretation.

4.4 Future work

Upcoming research should focus on experimental work. Since bone volume and

transducer frequency cannot be adjusted in experiments, and most vertebral specimens

are from older donors with lower bone volume, the major variable is trabecular bone

thickness. Therefore future experiments should include imaging of small cervical and

possibly upper thoracic pedicles to determine if consistent visibility of the cortical bone

can be attained. Furthermore, adjustment of the transmitted pulse, such as the pulse

energy, can be attempted to improve image quality. Finally, image post-processing such

as time-gain compensation can be utilized to compensate for the high attenuation, which

remains a primary obstruction to successful ultrasound imaging in bone.

71

References Anderson CC, Marutyan KR, Holland MR. Interference between wave modes may contribute to the apparent negative dispersion observed in cancellous bone. The Journal of the Acoustic Society of America. September 2008; 124:1781-1789. Bledsoe JM, Fenton D, Fogelson JL, Nottmeier EW. Accuracy of upper thoracic pedicle screw placement using three-dimensional image guidance. The Spine Journal. 2009; 9:817-821. Bossy E, Laugier P, Peyrin F, Padilla F. Attenuation in trabecular bone: A comparison between numerical simulation and experimental results in human femur. The Journal of the Acoustic Society of America. October 2007; 122:2469-2475. Chaffaî S, Peyrin F, Nuzzo S, Porcher R, Berger G, Laugier P. Ultrasonic characterization of human cancellous bone using transmission and backscatter measurements: relationships to density and microstructure. Bone. January 2002; 30:229-237. Cobbold R.S.C. (2007). Foundations of Biomedical Ultrasound. Oxford University Press, New York. Fu TS, Wong CB, Tsai TT, Liang YC, Chen LH, Chen WJ. Pedicle screw insertion: computed tomography versus fluoroscopic image guidance. International Orthopaedics. 2008; 32:517-521. Girardi FP, Cammisa FP, Sandhu HS, Alvarez L. The placement of lumbar pedicle screws using computerised stereotactic guidance. The Journal of Bone & Joint Surgery (Br). September 1999; 81-B:825-829. Hartl R, Theodore N, Dickman CA, Sonntag VKH. Technique of Thoracic Pedicle Screw Fixation for Trauma. Operative Techniques in Neurosurgery. March 2004; 7:22-30. Hosokawa A. Ultrasonic pulse waves in cancellous bone analyzed by finite-difference time-domain methods. Ultrasonics. 2006; 44:e227-e231. İnceoğlu S, Kilinçer C, Tami A, McLain RF. Cortex of the pedicle of the vertebral arch. Part II: microstructure. Journal of Neurosurgery: Spine. September 2007; 7:347-351. Ito Y, Sugimoto Y, Tomioka M, Hasegawa Y, Nakago K, Yagata Y. Clinical accuracy of 3D fluoroscopy-assisted cervical pedicle screw insertion. Journal of Neurosurgery: Spine. November 2008; 9:450-453. Kaufman JJ, Luo G, Siffert RS. Ultrasound simulation in bone. IEEE Transactions on Ultrasonics, Ferroelectrics, and Frequency Control. June 2008; 55:1205-1217.

72

Kim CW, Lee YP, Taylor W, Oygar A, Kim WK. Use of navigation-assisted fluoroscopy to decrease radiation exposure during minimally invasive spine surgery. The Spine Journal. 2008; 8:584-590. Kotil K, Bilge T. Accuracy of pedicle and mass screw placement in the spine without using fluoroscopy: a prospective clinical study. The Spine Journal. 2008; 8:591-596. Luo G, Kaufman JJ, Chiabrera A, Bianco B, Kinney JH, Haupt D, Ryaby JT, Siffert RS. Computational methods for ultrasound bone assessment. Ultrasound in Medicine & Biology. 1999; 25:823-830 Mandigo CE, Sampath P, Kaiser MG. Posterior Dynamic Stabilization of the Lumbar Spine: Pedicle-Based Stabilization with the AccuFlex Rod System. Neurosurgery Focus. January 2007; 22. Mano I, Yamamoto T, Hagino H, Teshima R, Takada M, Tsujimoto T, Otani T. Ultrasonic transmission characteristics of in vitro human cancellous bone. Japanese Journal of Applied Physics. 2007; 46:4858-4861. Moore KL, Dalley AF (2006). Clinically Oriented Anatomy – Fifth Edition. Lippincott Williams & Wilkins, Baltimore. Mujagic M. Ultrasound Guided Placement of Pedicle Screws for Spinal Fusion Surgery. University of Toronto MASc thesis, 2007. Onose G, Anghelescu A, Muresanu DF, Padure L, Haras MA, Chendreanu CO, Onose LV, Mirea A, Ciurea AV, El Masri WS, von Wild KRH. A review of published reports on neuroprotection in spinal cord injury. Spinal Cord. 2009; 47:716-726. Padilla F, Bossy E, Haïat G, Jenson F, Laugier P. Numerical simulation of ultrasound transmission in cancellous bone. IEEE Ultrasonics Symposium. 2005; 4:2022-2025. Vaccaro AR (2005). Spine – Core Knowledge in Orthopaedics. Elsevier Mosby, Philadelphia. Wear K. A model to predict dispersion in cancellous bone. IEEE Ultrasonics Symposium. 2001; 1217-1220. Zindrick MR, Wiltse LL, Doornik A, Widell EH, Knight GW, Patwardhan AG, Thomas JC, Rothman SL, Fields BT. Analysis of the morphometric characteristics of the thoracic and lumbar pedicles. Spine. March 1987; 12:160-166.