Ultrasound-Guided Chemodenervation Procedures

8. UPPER MOTOR NEURON SYNDROMES
41
Figure 8.6 Post-stroke spastic-
ity, motor impairments affect-
ing gait.
(a)
(b)
(c)
Figure 8.7 Casting and splinting for upper motor neuron syndrome (a) cerebral palsy, (b) stroke, and
(c) traumatic brain injury.
Chemodenervation Procedures for Muscle Hypertonia Associated with
UMNS—Patient Selection
In many clinical practices, muscle hypertonia associated with UMNS is the most com-
mon reason for patient referral for botulinum neurotoxin (BoNT) injections or other
chemodenervation procedures. Prior to any treatment, a detailed assessment of the
patient’s sensorimotor impairments, including hypertonia and function is required.
This assessment should include evaluation of tone and function and provides the basis
for patient selection, identifi cation of target muscles, as well as selection of the most
appropriate intervention(s). The assessment should determine the degree and severity

ULTRASOUND-GUIDED CHEMODENERVATION PROCEDURES
42
of the patient’s hypertonia and whether the hypertonia is problematic or symptomatic
(28–40). This evaluation should also identify realistic goals for chemodenervation ther-
apy. Because hypertonia is readily observed, patients, caregivers, and therapists often
report problematic hypertonia, muscle stiffness, pain, spasms, and limited range of
motion (passive and active), and that “tone” limits function or quality of life. These indi-
viduals, and sometimes their treating physicians, may fail to recognize the full impact
of various UMNS-related negative symptoms on the patient’s function. In a patient with
severe weakness or limited motor control, simply removing or reducing spasticity may
not improve active or passive function. This detailed patient assessment will identify all
of the patient’s sensorimotor impairments and guide goal setting and treatment selec-
tion (18,19,24,25).
Although physicians most frequently recommend chemodenervation procedures
(BoNT, nerve or motor point blocks for management of hypertonia associated with
UMNS), the decision to treat symptomatic hypertonia should also involve the patient
and others involved in their care. Muscle selection is critical and should be determined
by the history, information provided by caregivers or therapists, physical examination,
and functional assessment whereby the patient is observed carrying out various ADLs.
BoNT Therapy for Muscle Hypertonia
There are currently three type A and one type B botulinum toxins approved by the Food
and Drug Administration (FDA) for a variety of clinical uses in the United States (22–25).
Only onabotulinumtoxinA currently carries an indication in the United States of America
(USA) for treatment of spasticity (upper limb) (41–44), “Off label” uses for lower extremity
spasticity, and pediatric UMNS are permitted in the USA and such indications for BoNT
therapy are “on label” in Europe and other countries. A full discussion of current FDA
indications for botulinum toxin products currently available in the United States is pro-
vided elsewhere in this text.
The clinical effects of BoNTs injected in muscle is a dose dependent reduction in
muscle activity and tone. Toxin dosing is product specifi c and detailed instructions are
available in the respective product labels (see toxin basics chapters). The dose of toxin
administered is determined by the disorder being treated. For example, patients with
spasticity related to UMNS generally require higher doses of BoNTs than that required
for treating patients with focal dystonias. The responsiveness of an individual patient to
BoNT treatment is also quite variable. Therefore, in toxin naïve patients, the initial dose
should be carefully selected. The dose may then be incrementally increased over a series
of treatment sessions with the goal of using the lowest dose possible that provides mus-
cle relaxation suffi cient to achieve the functional goals set during treatment planning.
Larger doses of BoNT are, however, more commonly used in patients with severe spas-
ticity, in large muscles and when passive functional goals are selected. In some patients,
with severe spasticity or dystonia, physicians may recommend combined therapy with
BoNT and phenol (or alcohol) nerve or motor point blocks. Phenol can be used in a
“toxin sparing” technique. This may be required when the calculated dose of toxin to
treat the identifi ed target muscles exceeds recommended dose limits. For example, in
a patient with severe lower extremity spasticity, adductor tone may be addressed with
phenol obturator nerve blocks, “saving” toxin for other muscles (18–20, 35).
The clinical effects of BoNT therapy generally last 3 months but the benefi ts of treat-
ment may last longer in some patients (up to 6 months). The effects of phenol and alcohol
chemodenervation are longer lasting with reported benefi ts from 6 to 18 months. This

8. UPPER MOTOR NEURON SYNDROMES
43
extended benefi t is due to the protein denaturation/Wallerian degeneration produced by
phenol/alcohol chemodenervation procedures. Many patients will benefi t from repeat
BoNT treatment at 3 to 6 month intervals. Reevaluation of the patient, the extent and pat-
tern of muscle hypertonia should be performed prior to each treatment cycle.
Treatment Following Chemodenervation
Simply removing spasticity or muscle hypertonia may not improve function (19,20).
Following chemodenervation many, if not most, patients benefi t from additional therapy
(19–21,26,27). This may include increased passive stretching, splinting, and positioning
aids to improve range of motion. Therapy directed at improving active functional tasks
may include direct physical, occupational, and or speech therapy. Prior to chemoden-
ervation procedures, the need for and access to additional therapy should be discussed
with the patient and caregiver.
SPECIAL POPULATIONS
Pediatric patients represent a special subgroup presenting with UMNS as it relates to BoNT
and other chemodenervation procedures. Chemodenervation with phenol or alcohol has
been used for many decades, and BoNT therapy for nearly 2 decades (35,45,46). Treatment
goals are similar to those for adults and include improved passive and active function,
brace/splint tolerance, and reduced contracture/deformity (27,41) (Figure 8.8). Additional
goals include reducing post surgical pain and muscle spasm (47–49). The dose for toxin
therapy for children is generally calculated by body weight (27,41,42,44). Reported total
dose limits vary signifi cantly among individual practitioners. There are risks with high
dose protocols including generalized or local weakness and reported dysphagia. Reports of
aspiration events and death following BoNT therapy led the FDA to add a boxed warning
Figure 8.8 Gait dysfunction.

ULTRASOUND-GUIDED CHEMODENERVATION PROCEDURES
44
for all BoNT used in the United States. It is important to recognize that the FDA did not
conclude that toxin therapy was the cause of death in these patients.
SUMMARY
Many patients with UMNS-related muscle overactivity will benefi t from chemoden-
ervation procedures including BoNT and or nerve/motor point blocks. The success of
these procedures is dependent on identifying patients who may benefi t from treatment,
setting realistic goals, selecting the most appropriate treatment modality for a patient,
and correctly targeting problematic muscles.
REFERENCES
1. Lance, JW. Pyramidal and extrapyramidal disorders. In: Shani DT, ed. Electromyography in cns Disorders. Boston, MA:
Butterworth; 1984:1–19.
2. Lance, JW. The control of muscle tone, refl exes, and movement: Robert Wartenberg lecture. Neurology.
1980;30(12):1303–1313.
3. Lance JW. Symposium synopsis. In: Feldman RG, Young RR, Kella WP eds. Spasticity: Disordered Motor Control.
Chicago, IL: Year Book Medical Publishers; 1980 485–494.
4. Messina C. Pathophysiology of muscle tone. Funct Neurol. 1990;5(3):217–223.
5. Sheean G, McGuire JR. Spastic hypertonia and movement disorders: pathophysiology, clinical presentation, and
quantifi cation. PM&R. 2009;1(9):827–833.
6. Burke D. Spasticity as an adaptation to pyramidal tract injury. Adv Neurol. 1988;47:401–423.
7. Kuypers HGJM: Anatomy of the descending pathways. In: Brooks VB ed, Handbook of Physiology. Bethesda, MD:
American Physiological Society 1981; 597–666.
8. Mayer N. Clinicophysiologic concepts of spasticity and motor dysfunction in adults with an upper motor neuron le-
sion. In: Simpson DM, Mayer NH eds, Spasticity: Etiology, Evaluation, Management and the Role of Botulinum Toxin. New
York, NY: We Move Self Study CME; 2002: 1–11.
9. Sherrington CS. Decerebrate rigidity and refl ex coordination of movements. J Physiol. 1898;22:319–322.
10. Sherrington CS. On plastic tonus and proprioceptive refl exes. Quart J Exp Physiol. 1909;2:109–156.
11. Sherrington CS. Remarks on some aspects of refl ex inhibition. Proc Roy Soc B. 1925;97:519–545.
12. Schwindt PC. Control of motor neuron output by pathways descending from the brainstem. In: Towe AL, Luschei ES
eds, Handbook of Behavioral Physiology vol 15, Motor coordination. New York, NY: Plenum Press; 1981:139–230.
13. Nathan P. Some comments on spasticity and rigidity. In: Desmedt JE ed, New Developments in Electromyography and
Clinical Neurophysiology. Basel, Karger, 1973:13–14.
14. Pandyan AD, Gregoric M, Barnes MP, et al. Spasticity: clinical perceptions, neurological realities and meaningful mea-
surement. Disabil Rehabil. 2005;27(1–2):2–6.
15. Lin FM, Sabbahi M. Correlation of spasticity with hyperactive stretch refl exes and motor dysfunction in hemiplegia.
Arch Phys Med Rehabil. 1999;80(5):526–530.
16. Herman R. The myotactic refl ex: clinicophysiological aspects of spasticity and contracture. Brain. 1970;93:272–312.
17. Herman R, Freedman W, Mees S. Physiological aspects of hemiplegic and paraplegia spasticity. In: Desmedt JE ed,
New Developments in Electromyography and Clinical Neurophysiology. Human Refl exes, Pathophysiology of Motor Systems,
Methodology of Human Refl exes. Basel, Karger, 1973:579–589.
18. Mayer N, Esquenazi A, Childers MK. Common patterns of clinical motor dysfunction. Muscle Nerve. 1997;20(suppl
6):S21-S35.
19. Mayer N. Clinicophysiologic concepts of spasticity and motor dysfunction in adults with an upper motor neuron
lesion. Muscle and Nerve Suppl. 1997;6:S1-S13.
20. Wissel J, Schelosky LD, Scott J, Christe W, Faiss JH, Mueller J. Early development of spasticity following stroke: a
prospective, observational trial. J Neurol. 2010;257(7):1067–1072.
21. Lundstrom E, Terent A, Borg J. Prevalence of disabling spasticity 1 year after fi rst-ever stroke. J Neuro.
2008;15:533–539.

8. UPPER MOTOR NEURON SYNDROMES
45
22. Lundström E, Smits A, Terént A, Borg J. Time-course and determinants of spasticity during the fi rst six months follow-
ing fi rst-ever stroke. J Rehabil Med. 2010;42(4):296–301.
23. Pandyan AD, Cameron M, Powell J, Stott DJ, Granat MH. Contractures in the post-stroke wrist: a pilot study of its time
course of development and its association with upper limb recovery. Clin Rehabil. 2003;17(1):88–95.
24. Damiano DL, Alter KE, Chambers H. New clinical and research trends in lower extremity management in ambulatory
children with cerebral palsy. Phys Med Rehabil Clin N Am. 2009;20(3):469–491.
25. Elovic EP, Esquenazi A, Alter KE, et al. Chemodnervation and nerveblocks in the diagnosis and management of spas-
ticity and muscle overactivity. PMR. 2009;1(9):842–851.
26. Alter KE, Childers MK, Cindy B. Ivanhoe CB. (Chairperson Localized Pharmacologic Therapeutic Options: Current
and Novel Treatments to Optimize Upper- and Lower-Limb Spasticity Outcomes CME/CE 06/11 Medscape CME
course.
27. Gormley ME, O’Brien CF, Yablon SA. A clinical overview of treatment decisions in the management of spasticity. In:
Spasticity: Etiology, Evaluation, Management and the Role of Botulinum Toxin. New York, NY: We Move; 2002, WeMove.org.
28. Ansari NN, Naghdi S, Hasson S, Azarsa MH, Azarnia S. The Modifi ed Tardieu Scale for the measurement of elbow
fl exor spasticity in adult patients with hemiplegia. Brain Injur. 2008;22(13–14):1007–1012.
29. Ansari NN, Naghdi S, Arab TK, Jalaie S. The interrater and intrarater reliability of the Modifi ed Ashworth Scale in the
assessment of muscle spasticity: Limb and muscle group effect. Neuro Rehabilitation. 2008;23:231–237.
30. Bhimani RH, Anderson LC, Henly SJ, Stoddard SA. Clinical measurement of limb spasticity in adults: state of the sci-
ence. J Neurosci Nurs. 2011;43(2):104–115.
31. Haas BM, Bergstrom E, Jamous A, Bennie, A. The inter rater reliability of the original and of the Modifi ed Ashworth
Scale for the assessment of spasticity in patients with spinal cord injury. Spinal Cord. 1996;34:560–564.
32. Morris S. Ashworth and Tardieu Scales: their clinical relevance for measuring spasticity in adult and paediatric neu-
rological populations. Phys Ther Rev. 2002;7:53–62.
33. Ashford S, Turner-Stokes L. Goal attainment for spasticity management using botulinum toxin. Physiother Res Int.
2006;11:24–34.
34. Collin C, Wade DT, Davies S, Horne V. The Barthel ADL Index: a reliability study. Int Disabil Stud. 1988;10(2):61–63.
35. Wood KE. The use of phenol as a neurolytic agent, a review. Pain. 1978;5:205–229.
36. Dromerick AW, Edwards DF, Diringer MN. Sensitivity to changes in disability after stroke: a comparison of four scales
useful in clinical trials. J Rehabil Res Dev. 2003 Jan-Feb;40(1):1–8
37. Fosang AL, Galea MP, McCoy AT, Reddihough DS, Story I. Measures of muscle and joint performance in the lower
limb of children with cerebral palsy. Dev Med Child Neurol. 2003 Oct;45(10):664–70.
38. Lang, C.E., Wagner, J.M, Dromerick, A.W., & Edwards, D.F. (2006). Measurement of upper extremity function early
after stroke: properties of the action research arm test. Archives Physical Medicine and Rehabilitation, 87, 1605–1610.
39. Mahoney FI, Barthel DW. Functional evaluation: The Barthel Index. Md State Med J. 1965;14:61–65.
40. Nijland R, van Wegen E, Verbunt J, van Wijk R, van Kordelaar J, Kwakkel G. A comparison of two validated tests
for upper limb function after stroke: The Wolf Motor Function Test and the Action Research Arm Test. J Rehab Med.
2010;42:694–696.
41. Botox (Package Insert/Prescribing Information), Allergan, Irvine California
42. Dysport (Package Insert/prescribing information), Ipsen/Tercica, Brisbane, CA
43. Zeomin (Package Insert/Prescribing Information). Merz Pharmaceuticals, Greensboro,NC
44. Myobloc (Package Insert/Prescribing Information), Solstice Neurosciences, Malvern PA
45. Dromerick AW, Edwards DF, Diringer MN. Sensitivity to changes in disability after stroke: a comparison of four scales
useful in clinical trials. J Rehabil Res Dev. 2003;40(1):1–8.
46. Chung CY, Chen CL, Kuen-Wong AM. Pharmatherapy of spasticity in children with cerebral palsy. J Formos Med Assoc.
2011;110(4):215–222.
47. Graham, HK. Botulinum toxin type A management of spasticity in the context of orthopaedic surgery for children
with spastic cerebral palsy. Eur J Neurol. 2001;8(suppl 5):S30-S39.
48. Barwood S, Baillieu C, Boyd R, et al. Analgesic effects of botulinum toxin A: a randomized, placebo controlled trial.
Dev Med Child Neurol. 2000;42(2):116–121.
49. Alter KE. High-frequency ultrasound guidance for neurotoxin injections. Phys Med Rehabil Clin N Am.
2010;21(3):607–630.

Trunk
Codrin Lungu and Mark Hallett
Hyperhidrosis is a relatively prevalent condition defi ned as sweating in excess of the
normal requirements to regulate body temperature (1). The sweating response is medi-
ated by the sympathetic nervous system, and acetylcholine is the main neurotransmitter
of the eccrine gland synapses. Given the ability of botulinum toxin (BoNT) to temporar-
ily inhibit the release of acetylcholine, it can be successfully applied to the treatment of
hyperhidrosis (2).
Several randomized studies have demonstrated safety and effi cacy of BoNT treat-
ment of hyperhidrosis (3), allowing a reduction in sweating of around 75% of more (4).
Patients are selected for treatment depending on severity, and after failure of other
topical therapies. Excessive sweating can be quantifi ed by the minor starch iodine test,
which relies on a change of color in contact with sweat (5). Secondary hyperhidrosis due
to treatable conditions needs to be ruled out before BoNT therapy. The areas targeted
depend on individual patients, and typically include the axillary, palmar, and plantar
regions. Injections are performed in 20 to 40 locations distributed over the affected area
using a 26 G to 30 G needle. In large studies, doses used for axillary injections were 50
to 75 Units onabotulinumtoxinA and 200 to 250 Units abobotulinumtoxinA per axilla;
50 to 200 Units onabotulinumtoxinA and 120 to 500 Units abobotulinumtoxinA per
palm; 50 to 100 Units onabotulinumtoxinA per foot for plantar injections (5). Since
the injections are all superfi cial, ultrasound is generally not required or used for these
injections.
Compared to muscular injections for dystonia and spasticity, the effects of BoNT
treatment of hyperhidrosis are much longer lasting, with an average effi cacy duration
of 6 to 9 months (3,6). Treatment can be continued safely for years with good patient
satisfaction.
REFERENCES
1. Strutton DR, Kowalski JW, Glaser DA, Stang PE. US prevalence of hyperhidrosis and impact on individuals with axil-
lary hyperhidrosis: results from a national survey. J Am Acad Dermatol. 2004;51(2):241–248.
2. Swartling C, Naver H, Pihl-Lundin I, Hagforsen E, Vahlquist A. Sweat gland morphology and periglandular inner-
vation in essential palmar hyperhidrosis before and after treatment with intradermal botulinum toxin. J Am Acad
Dermatol. 2004;51(5):739–745.
9
Hyperhidrosis

4. TRUNCAL DYSTONIA AND OTHER TRUNCAL DYSKINESIAS
47
3. Doft MA, Hardy KL, Ascherman JA. Treatment of hyperhidrosis with botulinum toxin. Aesthet Surg J.
2012;32(2):238–244.
4. Heckmann M, Ceballos-Baumann AO, Plewig G. Botulinum toxin A for axillary hyperhidrosis (excessive sweating). N
Engl J Med. 2001;344(7):488–493.
5. Glaser D, Naumann M. Chapter 25: Botulinum neurotoxin in the management of hyperhidrosis and other hypersecre-
tory disorders. In: Jankovic J, Albanese, A., Atassi, MZ et al. eds. Botulinum Toxin Therapeutic Clinical Practice and Science.
Philadelphia, PA: Elsevier; 2009.
6. Doft MA, Kasten JL, Ascherman JA. Treatment of axillary hyperhidrosis with botulinum toxin: a single surgeon’s expe-
rience with 53 consecutive patients. Aesthetic Plast Surg. 2011;35(6):1079–1086.

Trunk
Katharine E. Alter, Zachary Bohart, and Codrin Lungu
Sialorrhea refers to the involuntary escape of saliva from the mouth and is a com-
mon reason for which patients are referred for botulinum toxin (BoNT) therapy.
While sialorrhea occurs in a number of conditions, the majority of patients present-
ing for BoNT therapy for sialorrhea have neurological disorders, most commonly
cerebral palsy, stroke, Parkinson disease, and amyotrophic lateral sclerosis (ALS).
These conditions cause a variety of sensory/motor impairments including impaired
sensation, lip closure, chewing, and swallowing (1–4). Sialorrhea often causes skin
breakdown or maceration and can have a signifi cant impact on quality of life includ-
ing soiling of clothing and social isolation. In addition, dysphagia or poorly con-
trolled swallowing may contribute to aspiration of oral secretions and resultant
aspiration pneumonia.
In the vast majority of patients referred for BoNT injections, sialorrea is caused by
disordered swallowing, not by an increase in saliva production. Management of sialor-
rhea may include oral motor therapy, devices to remove saliva from the face or mouth,
enteral medications, BoNT injections, and surgical procedures such as salivary gland
duct rerouting (5–10). BoNT therapy is increasingly offered to patients for whom more
conservative treatments such as pharmacotherapy do not provide adequate control of
saliva or cause intolerable side effects. BoNT is now often recommended prior to con-
sidering surgical interventions (8,9,11).
NORMAL SALIVATION
Saliva is produced by the three pairs of major salivary glands, as well as minor salivary
glands, which are scattered throughout the oral cavity. The volume of saliva produced in
adults is approximately 1.3 L per day. Saliva is important in maintaining dental health,
for lubrication of food, swallowing, and for digestion (enzymatic action of amylase).
Seventy percent of saliva is produced by the submandibular glands, 30% by the parotid
glands, 5% by the sublingual glands, and 5% by other minor salivary glands. The sub-
mandibular glands secrete the majority of basal saliva secretion at rest, whereas parotid
gland secretion increases in response to eating and other stimulation such as olfactory
stimulation (l). The salivary glands secrete saliva in response to parasympathetic, cho-
linergic stimulation via the release of acetylcholine (1,2,10–13). Because BoNT acts to
selectively reduce the release of acetylcholine at both neuromuscular and neuroglandu-
lar junctions, it can be an effective agent to reduce saliva production and sialorrhea in
patients with neurological disorders.
10
Botulinum Toxin Therapy for the
Management of Sialorrhea

10. BOTULINUM TOXIN THERAPY FOR THE MANAGEMENT OF SIALORRHEA
49
ORAL MOTOR CONTROL AND SIALORRHEA
Drooling is normal in infants and young children. Minor drooling may continue in early
childhood and has been reported in children up to the age of 5 years. Drooling or sia-
lorrhea that persists beyond this time is generally considered pathologic and is most
often seen in patients with neurologic disorders, as noted above. Sialorrhea is often
exacerbated by dental malocclusion, poor lip closure/oral motor control and limited
head or trunk control associated with neurologic dysfunction. For example, sialorrhea
is reported to occur in as many as 10% to 37% of children with cerebral palsy (1,2,10,11).
When patients have poor lip closure, limited oral motor control, and reduced or dis-
ordered swallowing, saliva pools in the fl oor of the mouth and spills over the lower
lip onto the chin, clothing, electronic equipment such as augmentative communication
devices, or wheelchair controls causing these devices to malfunction. Management of
sialorrhea is recommended when the degree and severity of sialorrhea poses such clin-
ical problems for a patient.
BOTULINUM TOXIN THERAPY FOR SIALORRHEA
BoTNs inhibit the release of acetylcholine at the salivary gland neuroglandular junc-
tions, thereby reducing saliva production. Injection of BoNT into salivary glands leads
to a dose dependent reduction in saliva production (14). There is a dearth of published
data on BoNT for sialorrhea related to the most appropriate starting dose and or the
maximum dose for the various BoNT products. RimabotulinumtoxinB, because of its
reportedly higher affi nity for autonomic nerves and incidence of anticholinergic side
effects, may have some theoretical advantage over type A toxins when treating patients
with sialorrhea but this theory is yet to be proven in head-to-head trials. Of note is the
report by Guidubaldi et al. (15) who found that duration of effect of type B toxin was
shorter than that of a type A serotype. The reported differences in duration of effect and
reduction of saliva production may be due to individual characteristics of type A and
B serotypes including dosing protocol, diffusion characteristics, and the affi nity of the
toxin serotype for autonomic nerve (15,16).
SIALORRHEA: REPORTED BOTULINUM TOXIN DOSE RANGES
Adults
Type A serotypes: In adults, the reported total mean toxin dose ranges from 55 to 300
Units of onabotulinumtoxinA and 250 to 450 Units of abobotulinumtoxinA. Starting
doses for onabotulinumtoxinA of 5 to 30 Units for each submandibular gland and 5 to 75
Units reported for individual parotid glands. Published dose ranges for rimabotulinum-
toxinB range from 1,500 to 5,000 Units. In one study, 3,000 Units of rimabotulinumtoxinB

ULTRASOUND-GUIDED CHEMODENERVATION PROCEDURES
50
was reported to be as effective as 5,000 Units with fewer side effects (15,17–19). For
rimabotulinumtoxinB, the reported dose range for individual salivary glands is 250 to
1000 Units for submandibular and 450 to 1000 Units for parotid glands (17). A com-
parison study reported similar effi cacy between abobotulinumtoxinA (250 Units) and
rimabotulinumtoxinB (2,500 Units) (15).
Pediatric Patients
In children, BoNT dose is often, but not always, calculated according to the patient’s body
weight in kilograms, rather than by the fi xed dose protocols used in adults. Units/Kg
dosing is used until the child achieves a weight when this dose is similar to the reported
dose used in adult protocols. Published toxin dose protocols in children range from 2 to
22.5 Units/Kg of onabotulinumtoxinA per gland. Reported dose ranges for rimabotu-
linumtoxinB in children range from 1,500 to 5,000 Units (total dose). In one study, no
added benefi t was noted and increased adverse events were reported when the dose
was increased from 1,500 Units to 3,000 Units of rimabolulinumtoxinB (16,17,19,20).
At Risk Populations: Patients With Motor Neuron Disease
Many patients with motor neuron disease, including patients with ALS, present with
symptomatic sialorrhea and may therefore benefi t from treatment with BoNT. Patients
with ALS are often fragile and have a degenerative condition leading to progressive
loss of motor function including swallowing and other bulbar motor functions. When
compared to sialorrhea caused by other neurological disorders, the majority of the
studies in patients with ALS report a similar effi cacy and side effect profi le for patients
receiving BoNT therapy for sialorrhea (4,11,22–24). However, a single case report has
been published documenting the acute deterioration of bulbar function in a patient
with ALS following BoNT injection (25). In light of the progressive nature of ALS and
the potential risk of weakness and dysphagia associated with BoNT therapy, clinicians
may choose a conservative approach to BoNT dosing when initiating BoNT therapy in
patients with ALS.
ACCURATE TARGETING OF SALIVARY GLANDS
When injecting BoNT into salivary glands of patients with sialorrhea, accurately tar-
geting these glands is important for effi cacy, as well as for reducing unwanted effects
of the procedure. Inaccurate targeting may lead to inadvertent injection of BoNT into
muscles that are directly adjacent to salivary glands. These muscles include the masse-
ter, pterygoids, and genioglossus. Inadvertent injection of these muscles may exacerbate
pre-existing oral motor problems, diffi culty with chewing, lip closure and dysphagia.
Electromyography (EMG) and electrical stimulation are not directly useful for local-
ization of salivary glands because although the absence of an audible EMG signal or
muscle twitch may indicate that the needle is not in an actively contracting muscle,
this does not indicate that the needle is in the desired gland. There are also signifi cant
limitations to anatomical guidance and palpation of salivary glands because of differ-
ences in patient size, anatomic variations, as well as the depth of the glands due to

10. BOTULINUM TOXIN THERAPY FOR THE MANAGEMENT OF SIALORRHEA
51
overlying subcutaneous tissue. Other than anatomic localization guides (12,15), high
frequency ultrasound (US) is the only localization technique which is able to increase
the accuracy of salivary gland localization when performing BoNT injections (20,26–29).
Salivary glands are easily distinguished from the surrounding muscle by the differ-
ences in echotexture between muscles and glands. Salivary glands, like other glands in
humans, have a uniform echogenic appearance on US scanning. This is distinctly differ-
ent from the mixed echotexture pattern seen in adjacent muscles (Figures 10.1a, 1b and
10.2a, 2b). For further detail, see next chapter on “Ultrasound Properties of Tissues.”
Studies indicate using US to guide BoNT injections of salivary glands leads to a greater
reduction in saliva fl ow and sialorrhea than blind injections, as well as a reduced risk
of adverse events. For a full discussion on salivary gland localization techniques, see
Chapter 16—“Ultrasound-Guided Botulinum Toxins: Clinical Applications.”
(a)
(b)
Figure 10.1 Parotid Gland Injection, Out of Plane Approach.
(a)
(b)
Figure 10.2 Submandibular Gland Injection, Out of Plane Approach.

ULTRASOUND-GUIDED CHEMODENERVATION PROCEDURES
52
REFERENCES
1. Harris SR, Purdy AH. Drooling and its management in cerebral palsy. Dev Med Child Neurol. 1987;25:804–814.
2. Tahmassebi JF, Curzon MEJ. The cause of drooling in children with cerebral palsy—hypersalivation or swallowing
defect? Int J Paediatr Dent. 2003;13(2):106–111.
3. Seppi K, Weintraub D, Coelho M, et al. The Movement Disorder Society evidence-based medicine review update:
Treatments for the non-motor symptoms of Parkinson’s disease. Mov Disord. 2011;26(suppl 3):S42-S80.
4. Blackhall LJ. Amyotrophic lateral sclerosis and palliative care: where we are, and the road ahead. Muscle Nerve.
2012;45(3):311–318.
5. Eiland LS.Glycopyrrolate for chronic drooling in children. Clin Ther. 2012;34(4):735–742.
6. Silvestre-Rangil J, Silvestre FJ, Puente-Sandoval A, et al. Clinical-therapeutic management of drooling: review and
update. Med Oral Patol Oral Cir Bucal. 2011;16(6):e763-e766.
7. Usta MG, Tufan AE, Cücelo lu EAJ Diphenhydramine use in the treatment of risperidone-induced sialorrhea [pub-
lished online ahead of print April 26, 2012]. Child Adolesc Psychopharmacol.
8. Intiso D. Therapeutic use of botulinum toxin in neurorehabilitation. J Toxicol. 2012;2012:802893.
9. Levy NS, Lowenthal DT. Application of botulinum toxin to clinical therapy: advances and cautions [published online
ahead of print March 24, 2012]. Am J Ther. .
10. Hornibrook J, Cochrane N. Contemporary surgical management of severe sialorrhea in children. ISRN Pediatr.
2012;2012:364875.
11. Ellies M, Laskawi R, Rohrbach-Volland S, et al. Up to date report of botulinum toxin therapy in patients with drooling
caused by different etiologies. J Oral Maxillofac Surg. 2003;61:454–457.
12. Rodriguez-Murphy E, Marti-Bonmati E, Camps-Seguí E, Bagán JV. Manually guided botulinum toxin type A subman-
dibular injections for the treatment of sialorrhea in tube-fed patients with advanced amyotrophic lateral sclerosis. Am
J Health Syst Pharm. 2011;68(18):1680–1681.
13. Lee JH, Lee BN, Kwon SO, et al. Anatomical localization of submandibular gland for botulinum toxin injection. Surg
Radiol Anat. 2010;32:945–949.
14. Chinnapongse R, Gullo K, Nemeth P, et al. Safety and effi cacy of botulinum toxin type B for treatment of sialorrhea in
Parkinson’s disease: a prospective double-blind trial. Mov Disord. 2012;27(2):219–226.
15. Guidubaldi A, Fasano A, Ialongo T, et al. Botulinum toxin A versus B in sialorrhea: a prospective, randomized, dou-
ble-blind, crossover pilot study in patients with amyotrophic lateral sclerosis or Parkinson’s disease. Mov Disord.
2011;26(2):313–319.
16. Intiso D. Therapeutic use of botulinum toxin in neurorehabilitation [published online ahead of print September 14,
2011]. J Toxicol.
17. Reddihough D, Erasmus CE, Johnson H, McKellar GMW, Jongerius PH. Botulinum toxin assessment, interven-
tion and aftercare for paediatric and adult drooling: international consensus statement. Eur J Neurol. 2010;17:(suppl
2):109–121.
18. Ondo WG, Hunter C, Moore W. A double-blind placebo-controlled trial of botulinum toxin B for sialorrhea in
Parkinson’s disease. Neurology. 2004;62:37–44.
19. Basciani M, Di Rienzo F, Fontana A, Copetti M, Pellegrini F, Intiso D. Botulinum toxin type B for sialorrhea in children
with cerebral palsy: a randomized trial comparing three doses. Dev Med Child Neurol. 2011;53(6):559–564.
20. Wu KP, Ke JY, Chen CY, Chen CL, Chou MY, Pei YC. Botulinum toxin type A on oral health in treating sialor-
rhea in children with cerebral palsy: a randomized, double-blind, placebo-controlled study. Child Neurol.
2011;26(7):838–843.
21. Tilton A, Vargus-Adams J, Delgado MR. Pharmacologic treatment of spasticity in children. Semin Pediatr Neurol.
2010;17(4):261–267.
22. Young CA, Ellis C, Johnson J, Sathasivam S, Pih N. Treatment for sialorrhea (excessive saliva) in people with motor
neuron disease/amyotrophic lateral sclerosis. Cochrane Database Syst Rev. 2011;(5):CD006981.
23. Jackson CE, Gronseth G, Rosenfeld J, et al. Randomized double-blind study of botulinum toxin type B for sialorrhea
in ALS patients. Muscle Nerve. 2009;39(2):137–143.
24. Gilio F, Iacovelli E, Frasca V, et al. Botulinum toxin type A for the treatment of sialorrhea in amyotrophic lateral scle-
rosis: a clinical and neurophysiological study. Amyotroph Lateral Scler. 2010;11(4):359–363.

10. BOTULINUM TOXIN THERAPY FOR THE MANAGEMENT OF SIALORRHEA
53
25. Meijer JW, van Kuijk AA, Geurts AC, et al. Acute deterioration of bulbar function after botulinum toxin treatment for
sialorrhea in amyotrophic lateral sclerosis. Am J Phys Med Rehabil. 2008;87(4):321–324.
26. Lee JH, Lee BN, Kwon SO, et al. Anatomical localization of submandibular gland for botulinum toxin injection. Surg
Radiol Anat. 2010;32(10):945–949.
27. Ellies, M, Laskawi R, Rohrbach-Volland S, et al. Botulinum toxin to reduce saliva fl ow: selected indications for ultra-
sound-guided toxin application into salivary glands. Laryngoscope. 2002;112(1):82–86.
28. Breheret R, Bizon A, Jeufroy C, et al. Ultrasound-guided botulinum toxin injections for treatment of drooling. Eur Ann
Otorhinolaryngol Head Neck Dis. 2011;128(5):224–229.
29. Dogu O, Apaydin D, Sevim S, et al. Ultrasound-guided versus ‘blind’ intraparotid injections of botulinum toxin-A for
the treatment of sialorrhea in patients with Parkinson’s disease. Clin Neurol Neurosurg. 2004;106(2):93–96.

(Trang không có chữ)

PART III
Ultrasound Basics
11. Physics of Ultrasound Imaging
12. Instrumentation and Knobology
13. Ultrasound for Procedural Guidance: Scanning—Techniques
and Tips
14. Chemodenervation: Ultrasound Characteristics of Relevant Tissues
15. Comparing Guidance Techniques for Chemodenervation Procedures
16. US-Guided Neurotoxin (BoNT) Injections: Clinical Applications
17. Ultrasound Guidance for Nerve and Motor Point Blocks

Trunk
Siddhartha Sikdar
Over the past 5 decades, ultrasound imaging technology has seen dramatic advances
that have established it as the diagnostic modality of choice for a wide variety of clini-
cal indications. Ultrasound provides real-time imaging capability that makes it an ideal
choice for image guidance during interventional procedures and for dynamic imag-
ing in a physician’s offi ce. Ultrasound energy at levels used for diagnostic imaging has
no known adverse effects on tissue, and the instrumentation is portable and relatively
inexpensive compared to other imaging modalities such as magnetic resonance imag-
ing (MRI) and computed tomography (CT) imaging. In recent years, advances in com-
puting technology have enabled dramatic miniaturization of ultrasound instruments,
and laptop-based ultrasound systems, hand-carried portable systems, and even pocket
ultrasound systems are now commercially available. With the advent of high-frequency
broadband ultrasound transducers, ultrasound imaging technology has become attrac-
tive for musculoskeletal applications, because it is able to visualize the structure of
muscles, tendons, and nerves, and diagnose abnormalities, including tears, fl uid accu-
mulation, infl ammation and fi brosis, in real time in the physician’s offi ce. This chapter
will review the basic physics of ultrasound and lay the foundation for an understanding
of the capabilities of ultrasound imaging technology, particularly for musculoskeletal
applications.
FORMATION OF AN ULTRASOUND IMAGE
Ultrasound imaging relies on the transmission of short bursts of sound into the human
body at high frequencies inaudible to the human ear. These sound waves are refl ected
back from interfaces between different tissue types, as well as scattered by the bulk soft
tissue microstructure. This refl ected and scattered sound energy is then processed to
produce the ultrasound image. The following sections will describe the basic physics of
sound waves.
Sound Waves and Their Properties
Sound is mechanical energy produced by the vibration of a source. This mechanical energy
can propagate from one location to another, and this propagating mechanical energy is
called a sound wave (also known as an acoustic wave). Sound energy requires an elastic
medium to propagate; sound does not travel in a vacuum. Acoustics is the fi eld of study
11
Physics of Ultrasound Imaging

11. PHYSICS OF ULTRASOUND IMAGING
57
that deals with the science of generation and propagation of mechanical waves. One way
to visualize the propagation of a type of mechanical wave is to imagine the ripples in a
lake when a pebble is dropped in the water. As the wave energy propagates, individual
particles in the medium move in place but do not propagate along with the wave. For
example, a leaf fl oating on the surface of the water continues to bobble up and down in
a to-and-fro motion while the wave passes by. This is called a transverse wave or a shear
wave because the particles of the medium (e.g., the leaf fl oating on water) are moving in a
direction perpendicular to the direction of the propagation of the wave. The acoustic energy
may also propagate as a compressional wave. One way to visualize a compressional wave
is to imagine a slinky. When one end of a slinky is moved back and forth, the mechanical
energy propagates along the slinky. The individual rings of the slinky move closer together
and then farther apart in a to-and-fro motion. In a compressional wave, the particles of the
medium are moving in the same direction as the propagation of the wave.
A fundamental aspect of this propagating mechanical energy is the vibration, or the
periodic to-and-fro oscillatory motion of the particles in the source that is producing
the sound and the particles in the medium through which the sound is propagating. To
understand the properties of the sound wave, we must fi rst understand some physical
properties associated with this propagating mechanical energy.
Frequency
Particles in the source of sound as well as particles in the medium through which
the sound is propagating oscillate to and fro. Frequency is defi ned as the number of
times the particles oscillate to and fro in 1 second. The unit of frequency is Hertz (Hz)
or cycles per second. Frequency is also referred to as pitch in the context of audible
sounds and music (e.g., middle C on a piano corresponds to a frequency of approx-
imately 256 Hz). The range of frequencies that are perceptible to the human ear is
20 Hz to 20,000 Hz. Frequencies above 20,000 Hz or 20 kilohertz (kHz) are referred to
as ultrasound. Typically for diagnostic ultrasound imaging, we use frequencies even
higher, typically 2 to 20 megahertz (MHz). The reason for using these high frequencies
in imaging will be apparent in the next sections.
Speed
The rate at which sound energy propagates through the medium is dependent on the
properties of the medium. Sound requires a medium to propagate and does not propa-
gate in a vacuum. Sound waves at ultrasonic frequencies propagate preferentially as com-
pressional waves, since the energy of the shear wave components are rapidly absorbed
by the medium. The speed of sound depends upon the density and compressibility of
the medium. Compressibility is an engineering concept that is related to the hardness or
softness of a medium. One way to measure this quantity is through the bulk modulus
of the medium, defi ned as the fractional change in the volume of the medium per unit
amount of uniform compressional pressure applied to the medium. The speed of sound
is given by c
K
=
ρ
, where c is the speed of sound in m/s, and ρ is the density of the
medium in kg/m3, and K is the bulk modulus. The unit of K is pascal (Pa) or kg/m/s2.

ULTRASOUND-GUIDED CHEMODENERVATION PROCEDURES
58
Sound travels faster in a medium that is stiffer and slower in a medium that is denser.
Therefore, the speed of sound in air (330 m/s) is lower than the speed in water and soft
tissue (~1,540 m/s), while the speed in bone is much higher (~3,400 m/s).
Wavelength
Another concept related to the propagation of sound is the wavelength. The wave-
length is defi ned as the distance between the peaks and troughs of the wave. The
propagating sound wave causes periodic peaks and troughs in the local pressure and
local density of the medium in addition to peaks and troughs in the local particle dis-
placement. The wavelength of sound is related to the speed and frequency as follows:
λ = c
f , where λ is the wavelength (in meters), c is the speed of sound (in m/s), and f is
the frequency (in cycles or Hz).
Therefore, the wavelength of a 10 MHz sound wave in water is
λ =
=
×
=
c
f
m/ s
Hz
mm
1540
10
10
1 54
1
× 0
0
m
0
0
−m =
154
6
4
54
1
× 0 m
The concept of wavelength is useful to understand why we use frequencies in the
ultrasonic range for diagnostic imaging. If the frequencies were lower, the wavelength
would be much larger than the typical size of scatterers in the human body. For exam-
ple, a 10 kHz sound wave has a wavelength of 0.15 m, which is signifi cantly larger than
most organs.
Intensity
The intensity of sound refers to the power (or energy per unit time) of the propa-
gating wave per unit area of the receiver surface. Intensity is typically measured in
Watts/cm2. Consider a point source of sound with power P. The sound wave propagates
as a spherical wave radiating outward from this source of sound. Now consider a spher-
ical shell at a distance r from the source of the sound. If the sound energy at this distance
is equally distributed in space, the amount of energy passing through a unit area of the
sphere, that is, the intensity is
P
r
4
2
π
, where 4πr2 is the area of the spherical shell at a
distance r. Therefore, intensity decreases with the reciprocal of the square of the dis-
tance as one moves away from a source. This is why sound is loud near the source of
the sound, but decreases rapidly farther away from the source. Intensity is an important
measure of an ultrasound wave, since it relates to safety and bioeffects. The bioeffects of
ultrasound will be discussed later in this chapter.
Sound Propagation Through a Medium as a Compressional Wave
Mechanical energy produced by the vibration of a source may propagate through a
medium either as a compressional wave or a transverse (shear) wave. At the high fre-
quencies associated with diagnostic ultrasound, the shear wave attenuates rapidly, thus
ultrasound energy is propagated predominantly as a compressional wave. As the com-
pressional wave passes through a medium, the values of the local density and local

11. PHYSICS OF ULTRASOUND IMAGING
59
pressure are perturbed from their equilibrium resting state and oscillate back and forth
(Figure 11.1). Therefore, there are local peaks followed by troughs in the pressure, den-
sity, as well as local particle velocity. A local peak positive pressure causes compression
of the medium and a local trough (i.e., peak negative) pressure causes rarefaction of
the medium. The local change in density affects the local speed of sound, since speed is
inversely proportional to the square root of density. This phenomenon causes nonlin-
earities in the propagation of the sound wave. At the instant of peak positive pressure,
the medium is denser, and the speed of sound is slower, while at the instant of peak
negative pressure, the medium is less dense and the speed of sound is faster. Therefore
the trough of the pressure wave propagates faster than the peak of the wave. These
nonlinearities can be exploited through ultrasound imaging. Many modern ultrasound
systems offer a harmonic imaging mode, that utilizes the higher frequencies generated
due to the nonlinear propagation of sound to create higher resolution images with fewer
imaging artifacts.
Formation of Ultrasound Beams
Classical wave theory, as formulated by the Dutch physicist Christiaan Huygens and
the French engineer Augustin-Jean Fresnel, provides a basis for understanding the spa-
tial distribution of energy as a sound wave, generated by an arbitrarily shaped source,
travels through an inhomogeneous medium. The sound wave from a point source of
sound radiates from this source as a spherical wavefront. The Huygens–Fresnel principle
Figure 11.1 Figure showing schematic of
compressional wave.
The compressional (or longitudinal) wave pass-
ing through a bulk medium causes particles
in the medium to alternatively move closer
together or farther apart, creating patterns of
compression and rarefaction. The local pres-
sure, density, and particle velocities oscillate to
and fro both in space and in time as the energy
propagates through the medium.

ULTRASOUND-GUIDED CHEMODENERVATION PROCEDURES
60
considers that any complex propagating wavefront can be considered as a collection
of numerous infi nitesimally small point sources generating new spherical wavefronts
(Figure 11.2). The location of the propagating wavefront at a subsequent instant in time
can be obtained by combining all these numerous spherical wavefronts. Recall that the
wave consists of peaks and troughs. Thus, combination of these wavefronts can lead
to positive enhancement when a peak combines with a peak, a negative enhancement
when a trough combines with a trough, and a cancellation when a peak combines with
a trough. This pattern of enhancement and cancellation leads to the diffraction of the
wave as it propagates. In the case of sound, the diffraction pattern depends upon the
geometry of the transducer producing the sound wave, as well as the structure of the
medium through which the sound propagates (1). For example, if the source of sound is
a piston, then sound energy is preferentially distributed as a beam in front of the piston
source, as shown in Figure 11.3.
The Huygens–Fresnel principle can be used to explain refl ection and refraction of
light, and the focusing of light using an optical lens. Sound energy may be focused in
the same way that light may be focused. In subsequent sections, we will describe how
array transducers can be used to focus sound waves.
Pulsed Echo Ultrasound Imaging
In diagnostic ultrasound imaging, a short pulse of sound is generated by an ultra-
sound transducer. The transducer is made of a material known as a piezoelectric, that
Figure 11.2 Figure schematically dem-
onstrating Huygens–Fresnel principle.
A propagating wavefront can be thought to
be composed of a number of point wavelet
sources each creating a spherical wavefront.
The propagating wavefront is then a superpo-
sition of these multiple spherical wavefronts.
The schematic shows a spherical wavefront
emanating from a point source, and a plane
wavefront emanating from a rectangular
source.

11. PHYSICS OF ULTRASOUND IMAGING
61
converts an electrical signal to mechanical energy and vice versa. The sound wave
propagates away from the transducer into the tissue as an ultrasound beam. As the
wave propagates, it encounters different tissue types and structures. A portion of the
sound energy is scattered and/or refl ected back to the transducer. The transducer
then converts this refl ected sound energy into an electrical signal that is processed by
the ultrasound system to form an image. This is the basis of pulsed echo ultrasound
imaging. The ultrasound imaging system measures the time duration between the
instant the ultrasound pulse is sent from the transducer, and the instant an echo is
received. This time duration is then converted to a measure of distance by assuming
a constant speed of sound. The strength of the refl ection is converted to brightness
in the image. Therefore an image of location of different tissue structures refl ecting
sound can be created, with the brightness corresponding to the strength of the refl ec-
tion from the tissue structures, which is the grayscale ultrasound image. The ability
to determine the location of a refl ecting tissue structure is known as ranging. In pulse
echo ultrasound, the accuracy of ranging is dependent upon the deviation of the
true speed of sound through the medium and the assumed nominal speed of sound
in the medium. Figure 11.4 shows how time duration between the transmission and
Figure 11.3 Figure showing the schematic beam profi le of a fl at piston transducer.
A compressional phase of the wave combining with another compressional phase leads to constructive interfer-
ence, whereas a compressional phase combining with a rarefactional phase leads to destructive interference. The
pattern of constructive and destructive interference leads to the beam profi le.

ULTRASOUND-GUIDED CHEMODENERVATION PROCEDURES
62
reception can be converted to distance. In subsequent sections, we will see how the
sound energy interacts with the tissue and how the refl ected energy is converted to
an image.
INTERACTION OF SOUND WITH TISSUE
Three important properties need to be considered in order to understand how sound
waves interact with soft tissue. The fi rst is the concept of acoustic impedance that deter-
mines the refl ection and refraction of sound as it propagates through the medium. The
second is the concept of scattering and diffraction as sound propagates through a struc-
tured medium. The third is the attenuation of sound energy by its absorption by the
medium. In the following sections, we will explore how these three properties infl uence
the creation of ultrasound images.
Depth into the Tissue
Distance along
the Tissue
Ultrasound
system
Ultrasound
system
Ultrasound
system
Time (t)
Depth d= cxt/2
1
2
3
An ultrasound pulse is
sent inside the tissue, and
echoes from different
tissue interfaces are
received at different times
The process is repeated
by sending pulses along
different scanlines
Scanlines
The strength of the echo
is converted to brightness
to generate the image
A
B
C
D
A
B
C
D
A
B
C
D
Figure 11.4 Figure showing schematic of pulse echo imaging with ranging.
Each individual pulse of ultrasound propagating through the layered medium is refl ected/scattered at inter-
faces, and the refl ected echoes arrive at the transducer at different time instants. The depth of the tissue interface
can be estimated from a knowledge of the speed of sound and the arrival time of the echo. The ultrasound image
is created by converting the brightness of the echo to a pixel intensity on the screen at the appropriate spatial
location.

11. PHYSICS OF ULTRASOUND IMAGING
63
Acoustic Impedance and Its Dependence on
Density and Speed of Sound
In electromagnetics, resistance is defi ned as the opposition to the fl ow of current in the
presence of a potential difference. Typically resistance leads to a loss of energy in the
form of heat. An analogous concept can be defi ned for the transfer of sound energy from
one point to another in a medium. The driving force in this case is the oscillation of the
pressure fi eld, which is analogous to potential. The particles in the medium respond to
this driving force and the particle velocity is analogous to current. Therefore, acoustic
impedance is defi ned as the change in particle velocity in response to a change in the
pressure. It can be shown that the acoustic impedance is dependent upon the density
of the medium and the speed of sound in the medium, according to the following rela-
tionship: Z = ρ × c, where Z is the acoustic impedance (in Rayls), c is the density of the
medium (kg/m3), and c is the speed of sound (in m/s). The unit of impedance is Rayls,
which corresponds to kg/m2/s in SI units.
Refl ection and Refraction at Interfaces
When sound energy travels from one medium (medium 1) with acoustic impedance
Z1 to another medium (medium 2) of acoustic impedance, Z2, a portion of the energy
is refl ected back into medium 1 while the remaining energy propagates forward into
medium 2. The refl ected intensity (Irefl ected), and transmitted intensity (Itransmitted) as a frac-
tion of the incident intensity (Iincident) is given by:
I
I
Z
Z
Z
Z
I
I
reflfe ected
incident
transmitted
inciden
= ⎛
⎝⎜
⎛
⎝
⎞
⎠⎟
⎞
⎠
2
1
Z
2
1
Z
2
;
t
Z
Z
Z
Z
Z Z
=
−⎛
⎝⎜
⎛
⎝
⎞
⎠⎟
⎞
⎠
=
1
4
2
1
Z
2
1
Z
2
2
1
Z
2
1
2
(
)
Z
Z
2
1
Z
In this case, we assume normal incidence, that is, the refl ecting medium is oriented
perpendicular to the direction of propagation of the sound wave. We can observe that
the fraction of refl ected intensity is determined by the difference in acoustic impedances
between two media. If the difference in acoustic impedance is high, a signifi cant amount
of the incident intensity will be refl ected back, whereas if the acoustic impedances are
very close to each other (i.e., Z1~Z2), most of the intensity is transmitted to the second
medium (Irefl ected~0; Itransmitted~Iincident).
There are several implications of the acoustic impedance of different tissue types for
ultrasound imaging. Looking at the numerical value of acoustic impedances for soft tis-
sue compared to air and bone in Table 11.1, we notice that most soft tissues have very
similar values of acoustic impedance, which implies that most of the sound energy is
transmitted through soft tissue, with some refl ection at the interfaces between tissue
types. For example, at the interface of fat and muscle, the amount of refl ected sound is
I
I
Z
Z
Z
Z
reflfe ected
incident
muscle
faf t
muscle
fat
= ⎛
⎝⎜
⎛
⎝
⎞
⎠⎟
⎞
⎠
=
−
2
1 7
1
.
.
7
1 38
3
1 7
1 38
0 104
0 0108
1
2
2
.7
1
.
.
104
0
%
+
⎛
⎝
⎛
⎝
⎛
⎞
⎠
⎞
⎠
⎞
=
=
0 104
≈

ULTRASOUND-GUIDED CHEMODENERVATION PROCEDURES
64
However, there is a large difference in acoustic impedance between soft tissue and
air, and between soft tissue and bone. An important consequence is that sound does not
effi ciently couple across an air–tissue interface or a tissue–bone interface. To couple the
sound energy from the transducer to tissue, we need an aqueous gel to remove the air
interface between the transducer and skin surface. The gel has acoustic impedance very
similar to that of soft tissue, and the acoustic impedance of the transducer is matched to
that of soft tissue, therefore one can get effi cient coupling of sound energy into the body.
When an air bubble is present between the transducer and the skin, the energy does not
couple effi ciently and a shadow is seen.
Similarly, when sound encounters bone, almost 43% of the energy is refl ected back
and very little is transmitted through bone.
I
I
Z
Z
Z
Z
reflfe ected
incident
bone
tissue
bone
tissue
= ⎛
⎝⎜
⎛
⎝
⎞
⎠⎟
⎞
⎠
=
−
2
7 8
1 6
1 3
7 8
1 63
0 654
0 428
43
2
2
8
1
.
.
654
0
%
+
⎛
⎝
⎛
⎝
⎛
⎞
⎠
⎞
⎠
⎞
=
=
0 654
≈
Therefore the bony interfaces appear bright because most of the energy is refl ected
back, and the tissues deeper than the bony interface cannot be visualized because very
little ultrasound energy reaches them. Figure 11.5 shows the appearance of a bony
interface. Note the bright appearance of the bone, and the shadow beneath it.
In the discussion above, we assumed normal incidence, that is the direction of the
propagating sound wave is perpendicular to the tissue interface. However, as the
operator manipulates the transducer for free hand scanning, the sound energy may
not necessarily impact an interface between two tissue types at a perpendicular angle.
We next consider the situation when the sound energy impacts the medium at an inci-
dence angle, θ > 0. The incidence angle is defi ned as the angle between the direction
of propagation of the sound wave and the normal to the interface. Using Huygens–
Fresnel principle, we can derive that a portion of the sound energy will be refl ected
back into medium 1, while the remaining energy will be transmitted into medium 2,
but will propagate at an angle that is different from the incident angle depending upon
the speeds of sound in the two media. This phenomenon is the same as that for light,
Table 11.1 Acoustic impedance of different tissues
Material
Acoustic Impedance (M Rayls, i.e. 106 Rayls)
Air
0.0004
Fat
1.38
Water
1.48
Blood
1.51 to 1.61
Kidney
1.62
Spleen
1.64
Liver
1.65
Soft tissue (average)
1.63
Muscle (average)
1.7
Bone
3.8 to 7.8

11. PHYSICS OF ULTRASOUND IMAGING
65
and is shown in Figure 11.6. The incident angle is always equal to refl ected angle. The
relationships between the incident angle θi, the refracted angle θt, and that between
the refl ected and incident intensities are determined by the following equation:
sin
sin
c
θ
θ
θ
cos
iθ
refl
e ected
incident
i
t
cosθ
cos
c
c
I
I
z cosθ
z
( )
( )
θtθ
=
f
=
1
2
2 cos
i
cos
2
;
os
o θ
θ
cos
i
t
cos
θ
θ
θ
θ
cos
⎛
⎝⎜
⎛
⎝
⎞
⎠⎟
⎞
⎠
2
;
Figure 11.5 Ultrasound image showing
bony interface (open arrow) and shadow-
ing (outlined arrow).
θ i
θr
Medium 1, c1, Z1
Incident wave
Reflected wave
θ t
Medium 2, c2, Z2
Refracted transmitted
wave
Figure 11.6 Refraction and refl ection of ultra-
sound at an interface between two media with
differing speeds of sound.

ULTRASOUND-GUIDED CHEMODENERVATION PROCEDURES
66
Figure 11.7 Ultrasound images showing the effect of normal incidence and deviation from normal inci-
dence at a tissue interface.
Figure 11.8 Reverberation artifact from
needle.
There is an important implication of this principle for ultrasound imaging. It implies
that for normal incidence of the ultrasound wave to an interface, the refl ected wave trav-
els back to the transducer approximately along the same path. However, if the ultrasound
beam is not perpendicular to the tissue interface, then the energy is refl ected back at a
different angle and some of the refl ected energy may not reach the transducer. Therefore
for optimal visualization of tissue interfaces, it is essential that the scanning be performed
so that the tissue interface is nearly perpendicular to the ultrasound beam (Figure 11.7).
Often, if the sound wave encounters two interfaces that are both mirror-like, then the
sound energy bounces back and forth causing a reverberation artifact (Figure 11.8). This
phenomenon is similar in principle to the multiple refl ections seen when two optical mir-
rors are placed in front of each other.

11. PHYSICS OF ULTRASOUND IMAGING
67
Scattering and Its Dependence on Tissue Structure
In the previous section, we treated refl ection from interfaces that were large compared to
the wavelength of sound. Such interfaces are called specular (or mirror-like) refl ectors,
since the refl ected sound wave behaves as if it were perfectly refl ected from a mirrored
surface. There are many examples of such specular refl ectors in musculoskeletal ultra-
sound imaging. For example, refl ections from bone, fascia, and tendon sheaths behave
approximately as specular refl ections (Figure 11.9).
However, most interfaces in the body are not ideal mirror-like and have some degree
of roughness to it. The roughness is often in the order of the wavelength of the typical
ultrasound waves that are used for diagnostic imaging. When sound energy is incident
on such a rough surface, some of the energy is scattered in directions away from the
direction of insonation (Figure 11.10). This enables visualization of some tissue inter-
faces, such as fascial planes, even if they are not exactly perpendicular to the direction of
insonation. On the other hand, some anisotropic structures, such as the collagen fi bers
in tendons, can no longer be adequately visualized if the angle of insonation deviates
signifi cantly from the perpendicular. This might cause portions of the tendon to appear
hypoechoic since the sound energy from those portions is refl ected away from the trans-
ducer (Figure 11.11).
As ultrasound propagates through bulk soft tissue, a portion of the acoustic energy is
typically scattered in all directions. The microstructure of tissue, in particular the size,
concentration, and orientation of the scatterers affects the amount of energy scattered
back toward the transducer. The backscattering cross section is one way to quantify the
amount of energy being backscattered to the transducer by objects that are not specular
refl ectors. The scattering by tissue depends upon the scale of the object relative to the
wavelength of ultrasound. Table 11.2 qualitatively lists different types of scattering and
their relationship to the wavelenght of ultrasound. Recall that 10 MHz ultrasound has a
wavelength in soft tissue of approximately 0.154 mm. Objects that are much larger than
this wavelength behave as specular refl ectors. Objects that are much smaller than this
wavelength, such as red blood cells, behave as Rayleigh scatterers, where the scattering
strength is weak and is strongly dependent on frequency.
Figure 11.9 Ultrasound images showing
specular refl ections from bony interface
(bottom arrow) and from muscle fascia
(top two arrows).

ULTRASOUND-GUIDED CHEMODENERVATION PROCEDURES
68
The echogenicity (i.e., the brightness of the refl ected echoes) of different tissues are
determined by their structure and acoustic properties relative to the surrounding tissue.
Therefore, fl uids and blood appear dark in ultrasound images. Fatty infi ltration of tissue
and fi brosis increases the echogenicity because of local changes in the acoustic imped-
ance. The size of these inhomegenieties in the tissue also affects echogenicity.
The main acoustic scatterers in muscle tissue are the connective tissue sheaths
around fascicles that create interfaces with different acoustic properties. In tendons, the
Table 11.2 Scattering from tissue
Scale of structure (L)
relative to wavelength
of sound (λ)
Scattering
strength
Frequency
dependence
Examples in the human body
L >> λ
Strong
~f
Diaphragm, large blood vessel walls, Bone
surface, fascial plane, tendon sheath
L ~ λ
Moderate
Variable
Most soft tissue structures
L << λ
Weak
~f4
Red blood cells
Figure 11.10 Schematic of ultrasound scattering.
When a pulse of ultrasound strikes a smooth interface between two different tissue types, it is refl ected back at
an angle following the laws of refl ection. The refl ected energy does not reach the ultrasound transducer, and such
an interface is invisible to the transducer. If the interface is rough (i.e., has imperfections at the scale of the ultra-
sound wavelength), then the sound waves are scattered in all directions, and some fraction of the incident energy
is backscattered towards the transducer. When sound propagated through bulk tissue, the tissue microstructure
scatters the sound in all directions and creates an interference pattern when it reaches the transducer. This inter-
ference leads to the phenomenon of ultrasound speckle.

11. PHYSICS OF ULTRASOUND IMAGING
69
Figure 11.11 Hypoechoic region
observed in a tendon due to a change in
the angle of insonation of an anisotropic
structure.
Figure 11.12 Ultrasound images showing speckle in muscle; longitudinal (left panel) and transverse
views (right panel) have markedly different speckle patterns. In longitudinal views, muscle fascicles are
visible as linear structures resulting in a somewhat fi brillar pattern. In transverse views, the muscle fas-
cicles are visible as points in cross section, resulting in a pattern that is sometimes referred to as a "starry
night" pattern.
scattering occurs at the interfaces between collagen fi bers. Similarly in nerves, the scat-
tering occurs at the interfaces of nerve fi bers. These tissues have very different appear-
ance in longitudinal images, where the transducer is aligned along the fi bers, and in
transverse images where the fi bers are visualized in cross section. The echotexture of
musculoskeletal ultrasound images are shown in Figure 11.12.
Speckle Formation and Its Dependence on Tissue Microstructure
As sound waves propagate through tissue, it undergoes scattering and also diffrac-
tion. Diffraction determines the behavior of a wave as it passes around obstacles. When

ULTRASOUND-GUIDED CHEMODENERVATION PROCEDURES
70
sound travels through a collection of scatterers, a characteristic pattern of constructive
and destructive interference will be formed due to diffraction. This granular noisy pat-
tern is pervasive in ultrasound images and is called speckle. This speckle pattern is a
unique signature of the microstructure of the tissue and the arrangement and size of the
scatterers in the tissue. Different tissues have different speckle patterns. The patterns
are stable, and if the transducer is held stationary and the tissue is not moving signifi -
cantly, the speckle pattern remains stationary. When the tissue moves, the speckle pat-
tern moves along with it.
The echotexture of muscles, tendons, and nerves depends upon the microstructure
and organization of the underlying tissue. In longitudinal images of muscle, the speckle
pattern is dominated by the longitudinal organization of the muscle fascicles and fascial
planes. This pattern is very different from that of the subcutaneous tissue. On the other
hand, in transverse images of muscle, where the fascicles are being viewed in cross sec-
tion, the speckle pattern is quite different as seen in Figure 11.12. Tendons have a dense
packing of collagen fi bers, thus their echotexture is different from muscle, although it
has a similar fi brillar pattern. Nerves viewed in cross section have a characteristic hon-
eycombed appearance because of the organization and packing of the nerve fi bers.
Attenuation of Sound
As sound travels through tissue, some of the energy is absorbed and converted to
heat. This energy loss of the sound wave is called attenuation. One mechanism by
which this energy dissipation is thought to occur is viscous damping in tissue. The
viscosity of tissue causes a time lag between the applied force and the tissue response.
This implies some energy storage in the tissue, which is then dissipated as heat. Other
mechanisms have also been proposed for the attenuation of sound. Higher frequen-
cies of sound waves suffer larger viscous losses. Therefore, ultrasound attenuation is
strongly frequency dependent, and the frequency dependence is different for differ-
ent types of tissue. The attenuation manifests as an exponential decrease of the sound
intensity with distance of propagation into the tissue. The intensity, I, therefore is a
function of ultrasound frequency and distance travelled into the body, and can be
written as:
z
I e
f z
,f
)
(f
−
0
α
where I0 is the initial intensity of the sound beam, and α(f) is the frequency dependent
attenuation coeffi cient.
The relative attenuation on a decibel (dB) scale can be written as:
10
4 34
0 23
0
log
.4
10
(
( )
.
I
I
f z
( )
(
f(
I
z
dB
⎛
⎝
⎛
⎝
⎞
⎠
⎞
⎠
Therefore, one way to specify units for attenuation is dB/cm/MHz, assuming a lin-
ear relationship between ultrasound attenuation coeffi cient and tissue. The attenuation
coeffi cients for different tissue types are given in the Table 11.3 below.

11. PHYSICS OF ULTRASOUND IMAGING
71
Frequency Dependence of Attenuation and Dependence on Material Properties
As discussed in the previous section, ultrasound attenuation is strongly dependent on
frequency. The physics of acoustic propagation in soft tissue is not linear, and the fre-
quency dependence of the attenuation coeffi cient is often nonlinear. Measurements of
the attenuation coeffi cient at different frequencies in human tissue have been reported
in the literature (2). The frequency dependence of the attenuation coeffi cient is different
for different tissue types. The attenuation coeffi cient of muscle tissue does not change
much with frequency, while that of other soft tissues on average has a stronger depen-
dence on frequency. An important consequence of this phenomenon is that lower fre-
quencies of sound are attenuated less than higher frequencies. Therefore, imaging at a
lower frequency increases penetration and enables visualization of deeper structures
compared to a higher frequency (Figure 11.13).
Figure 11.13 Lower frequencies allow deeper penetration compared to higher frequencies.
Table 11.3 Attenuation coeffi cients of different materials
Material
Attenuation Coeffi cient (db/
cm/MHz) at 1 MHz
Water
0.0022
Blood
0.18
Fat
0.66
Soft tissue (average)
0.9
Muscle (average)
2.0
Air
12.0
Bone
20.0
Lung
40.0

ULTRASOUND-GUIDED CHEMODENERVATION PROCEDURES
72
Safety Considerations
There are currently no known adverse effects of ultrasound at intensities typically
used for diagnostic imaging (3). However, ultrasound energy at high intensity does
produce bioeffects in tissue, therefore users should be aware of the potential risks.
Sound is mechanical energy and at suffi ciently high intensities it can cause local heat-
ing of tissue. In some cases, if the peak negative pressure is too high, it can cause
small pockets of gas to form in the tissue, a process called cavitation. These bioeffects
are most critical for obstetric imaging, since developmental defects can occur in the
fetus if the rise in tissue temperature exceeds 4o C. These bioeffects are less critical
for routine musculoskeletal imaging, but operators should nevertheless follow the
ALARA (As Low As Reasonably Achievable) principle, and the acoustic output and
exposure time should be kept to a minimum for the diagnostic or interventional task.
The Food and Drug Administration (FDA) specifi es limits for the allowable intensity
and peak negative pressure of diagnostic instruments. In addition, the FDA Output
Display Standard (ODS) also requires manufacturers to display two indices on the
ultrasound instrument’s display: the thermal index (TI), which provides information
about the potential for tissue heating, and the mechanical index (MI), which provides
information about the potential for cavitation. The TI should preferably be kept below
3.0 (lower if imaging for an extended period of time), and the MI should be below 1.9
for all applications. The use of ultrasound contrast agents (gas fi lled microbubbles)
involves some additional safety considerations, and is currently not approved by the
FDA for applications other than echocardiography.
ULTRASOUND INSTRUMENTATION
Diagnostic ultrasound equipment was fi rst manufactured in the early 1960s. The fi rst
instruments involved large articulated arms with a single ultrasound transducer that
were physically moved to create static images. In the 1970s, real-time ultrasound
imagers were developed using an array of ultrasound transducers that were electron-
ically switched to produce images. Modern ultrasound instrumentation has made
rapid strides in the past decade. Premium ultrasound systems now have advanced
high frequency transducers and sophisticated signal processing capabilities to pro-
duce exquisitely detailed images. At the other end of the spectrum, technological
innovations in computing have enabled the miniaturization of ultrasound systems
and there are now several hand-held and portable ultrasound systems available
commercially.
One important component of an ultrasound system is the transducer that transmits
sound energy into the body and receives the backscattered echoes. The transducer con-
verts electrical signals to sound energy and also converts the received sound energy back
into electrical signals. The rest of the ultrasound equipment converts these electrical
signals into images. Most modern imaging systems utilize array transducers, that are
a composite of multiple smaller transducers, and most of the processing is nowadays
performed digitally. So, modern ultrasound systems are in essence powerful computers
running advanced software for signal and image reconstruction and processing and for
communicating with electronic circuitry that control the ultrasound transducers. In the
following sections, we will review the basic building blocks of an ultrasound imaging
system.

11. PHYSICS OF ULTRASOUND IMAGING
73
Ultrasound Transducers and Piezoelectric Materials
Ultrasound transducers convert electrical energy into mechanical energy and vice versa.
Materials that are capable of this energy conversion are called piezoelectric materials.
The unique characteristics that affect the choice of a suitable piezoelectric material for
a diagnostic ultrasound transducer is the ability to effi ciently convert electrical energy
to mechanical energy and vice versa with minimal losses to heat and uniformly high
sensitivity over a broad range of frequencies. The most common choice of piezoelectric
material for imaging transducers is a ceramic compound material, lead zirconate tita-
nate (PZT). Other materials, such as poly-vinyldine-difl uoride (PVDF), a fl exible copol-
ymer, is being explored in research applications.
Pulsed Echoes and Bandwidth
The key aspect of pulse echo imaging transducers is the ability to produce controlled
short duration bursts of ultrasound. The transducer design is geared toward achiev-
ing this while maintaining high sensitivity to detect weak received echoes. One way
to think of bandwidth of acoustic sources is to consider a tuning fork versus a wooden
table top. When a tuning fork is struck, it often keeps ringing for a prolonged period of
time, in other words, a short duration impulse elicits an extended oscillatory mechani-
cal response. Such a source is called a narrowband source. In contrast when one knocks
on a table top, it produces a dull thud. A short duration impulse elicits a short dura-
tion mechanical response. Such a source is called a broadband source. Narrowband
behavior is not desirable for ultrasound imaging transducers. A good ultrasound trans-
ducer when excited with a short duration electrical impulse should produce a very
short duration mechanical response to improve spatial resolution. The relationship
between spatial resolution and the duration of the ultrasound pulse will be described
later in this chapter. Therefore the mechanical response of the transducer needs to be
heavily damped. This is achieved by providing appropriate backing material to the
transducer. The frequency of the ultrasound transducer depends upon the width of the
PZT ceramic. According to the physics of acoustic waves, the thickness of the PZT ele-
ment is typically a half wavelength of the ultrasound waves.
Array Transducers and Focusing
Modern ultrasound systems utilize array transducers that comprise of multiple piezo-
electric elements that can be individually controlled electronically (1). A major advan-
tage of using an array transducer is the ability to focus the ultrasound beam at different
spatial locations by electronically controlling the elements of the transducer. An ultra-
sound beam focused at a specifi c location in space has increased acoustic energy at the
focus compared to other spatial locations. Focusing can be achieved by ensuring that
the ultrasound waves produced by the individual transducer elements in the array add
constructively in phase at the focus leading to an increase in the amplitude of the ultra-
sound pressure. For example, a peak compression phase of the pressure pulse combining
with another peak compression would lead to an increase in peak compressive pressure,
whereas a compression combining with a rarefaction would lead to destructive interfer-
ence and a decrease in pressure. Therefore, it is important to ensure that the individual

ULTRASOUND-GUIDED CHEMODENERVATION PROCEDURES
74
ultrasound pulses from each element in the array arrive at the same time at the focus of
the transducer. This can be accomplished by transmitting the pulses from each element
of the array at slightly different times, with precisely timed delays that are calculated
from the differences in path length between each element and the spatial location of
the focus. For example, Figure 11.14 shows the elements of an array transducer excited
at different times, such that the pulse arrive at a focal point in space at the same time
leading to an enhancement of the amplitude. This process is called transmit focusing or
transmit beamforming. A similar process can be performed during receiving the back-
scattered ultrasound echoes. The ultrasound energy scattered from a specifi c location in
space arrives at the different elements on the array transducer at slightly different times.
Therefore, if the received pressure pulses are combined without accounting for this time
delay, some of the signal is lost by destructive interference combining compressions with
rarefactions. By precisely delaying the signals received by each element before combin-
ing them, the strength of the signal can be enhanced. This is called receive focusing or
receive beamforming. Receive focusing is performed automatically by the ultrasound
instrument. However, the transmit focusing location can be manually set by the operator.
Figure 11.15 shows the effect of incorrectly setting the transmit focal zone. The locations
in the focal zone have good spatial resolution and contrast. If the focal zones were set too
deep, then the resolution and contrast becomes worse.
Figure 11.14 Focusing using an array transducer by precisely con-
trolling the timing of the transmission of each element. The transmis-
sions are delayed with respect to each other, so that all the individual
pulses arrive at the focal point at exactly the same time leading to a
constructive interference and enhanced signal amplitude.

11. PHYSICS OF ULTRASOUND IMAGING
75
Basic Processing Steps to Produce Ultrasound Images
The ultrasound processing system converts the backscattered echoes into pixels on an
image. The transducer receives the backscattered pressure wave, and the piezoelectric
elements convert this mechanical energy into electrical signals. These signals are then
processed by a beamformer, which performs receive focusing by combining the appro-
priately delayed received signals. At this point, the electrical signals corresponding to
the ultrasound echoes are in the radiofrequency (RF) range (several MHz), and are often
referred to as the RF signal. The strength of the backscattered echo modulates the ampli-
tude of this RF signal. In modern ultrasound systems, the electrical signals are then
converted into a digital form using high-speed analog to digital converters and much of
the processing is performed digitally. In some systems, beamforming is also performed
digitally. The next step is to detect the amplitude of the RF signal to extract the echo
strength. This processing step conceptually is very similar to an AM radio receiver that
demodulates the amplitude from a modulating RF carrier signal. Premium ultrasound
systems with Doppler capability utilize more sophisticated processing systems that pre-
serve information about the phase of the RF signal in addition to the amplitude. The
amplitude is mapped to the pixel brightness on the grayscale ultrasound image. To com-
pensate for the attenuation of ultrasound with depth, time gain compensation (TGC) is
applied that increases the amplifi cation of the signals from deeper tissue. The TGC can
be applied in hardware in the receiver preamplifi er, or in software. Most ultrasound
systems also have some image processing to improve the appearance of the image and
enhance the contrast. The basic processing steps are shown in Figure 11.16.
IMAGE RESOLUTION AND IMAGE CONTRAST
Two important diagnostic features of an ultrasound image are the spatial resolution that
determines the ability to discriminate between closely separated structures and better
defi ne small structures, and image contrast that determines the ability to discriminate
Figure 11.15 Images showing the effect of the transmit focal zone on the lateral image resolution. The
image on the left has a transmit focal zone (indicated by the bar and green arrows to the right of the
image) set at the appropriate depth of interest. The image on the right has the transmit focal zone set too
deep, and thus the imaging resolution of the muscle of interest is poorer.

ULTRASOUND-GUIDED CHEMODENERVATION PROCEDURES
76
between two different tissue types that are closely separated in terms of brightness. The
following sections will describe the underlying principles that determine the spatial res-
olution and image contrast.
Ultrasound Frequency and Its Relationship to Image Resolution
The spatial resolution of an imaging system is defi ned as the minimum physical separa-
tion between two point targets that are still distinguishable as separate. For optical sys-
tems, such as microscopes, this depends upon the quality of optics and focusing, but the
physics of light propagation still imposes some fundamental limits on the spatial reso-
lution, known as the diffraction limit. Similarly, for ultrasound imaging, there are some
fundamental limits on the spatial resolution imposed by the physics of acoustic wave
Scanlines
Beamformer
Scanlines
Demodulator/
Envelope detector
Image
processing/Filtering
Timing and
Control
electronics
Time Gain
Compensation
A
B
C
D
Depth into the Tissue
Distance along
the Tissue
Transmitter/Receiver
electronics
Display
Figure 11.16 Block diagram of the processing steps in an ultrasound system.

11. PHYSICS OF ULTRASOUND IMAGING
77
propagation. One important concept that is critical to understanding resolution in ultra-
sound images is the anisotropy of the ultrasound beam in the three spatial dimensions.
Figure 11.17 shows the geometry of an ultrasound beam in three dimensions showing
the different resolutions in each dimension. The axial direction is along the path of the
ultrasound beam and denotes the depth into the tissue, the lateral direction is perpen-
dicular to the axial direction in the imaging plane, while the elevational direction is per-
pendicular to the imaging plane. The image resolutions along these three dimensions
are different and depend upon different factors. The following sections describe the
relationship between image resolution and the frequency and focusing properties of the
ultrasound beam.
Axial Resolution and Its Dependence on Frequency and Bandwidth
The axial resolution is the minimum distance between two resolvable point targets along
the path of the ultrasound beam. If this minimum resolvable distance is small, the axial
resolution is high. Axial resolution is specifi ed in units of distance, mm or microns (µm),
for example, 0.1 mm=100 microns (µm). The axial resolution depends upon the spatial
extent of the ultrasound pulse. Figure 11.18 shows the relationship between the spatial
extent of the pulse and axial resolution. As can be seen from the schematic, the spatial
extent depends upon two factors: the wavelength of the ultrasound and the number of
cycles in the pulse. Recall from the section “Sound Waves and Their Properties” that
wavelength is inversely related to the frequency of the ultrasound wave. The wave-
length is small for higher frequencies, therefore the axial resolution is better for higher
frequencies.
The second factor that infl uences the axial resolution is the number of cycles in the
pulse. For a two-cycle pulse with wavelength 0.1 mm, the axial resolution is effec-
tively ~0.2 mm, while for a four-cycle pulse, it is effectively ~0.4 mm. Communication
Axial resolution
Elevational
resolution
Lateral
resolution
Figure 11.17 Three dimensional schematic of an ultra-
sound beam showing axial, lateral, and elevational spa-
tial resolution.

ULTRASOUND-GUIDED CHEMODENERVATION PROCEDURES
78
theory tells us that the number of cycles in a pulse is dependent upon the frequency
bandwidth of the pulse. A pulse with a smaller number of cycles has a larger fre-
quency bandwidth. High frequency broadband transducers are capable of producing
ultrasound pulses that have smaller spatial extent. Commercially available ultrasound
transducers are labeled in terms of the range of frequencies they can support. For
example, a Philips L12–5 transducer is capable of supporting frequencies between 5
MHz and 12 MHz, for a nominal bandwidth of 12 − 5 = 7 MHz. Compare that to a L7–4
transducer, which is capable of supporting frequencies between 4 MHz and 7 MHz,
with a nominal bandwidth of 7 − 4 = 3 MHz. The L12–5 transducer has a higher fre-
quency and a higher bandwidth, therefore is capable of better axial resolution than the
L7–4 transducer.
Lateral Resolution and Its Dependence on Focusing and Frequency
The lateral resolution is the minimum distance between two resolvable point targets along
the lateral direction. Similar to axial resolution, lateral resolution is also specifi ed in units of
distance. To be able to resolve two separate targets in the lateral direction, the ultrasound
echo from one target should be distinct from the echo from the other. If the ultrasound
beam is narrowly focused, then it can elicit distinct echoes from two closely separated
targets. Figure 11.19 shows a schematic of an ultrasound beam insonating two laterally
separated point targets, and the relationship between the width of the beam and the lateral
resolution. The width of an unfocused ultrasound beam, shown in Figure 11.3, is inversely
dependent on frequency. In addition, focusing can further narrow the width of the beam.
Therefore, the lateral resolution depends upon the frequency as well as focusing. The tight-
ness of the focusing depends upon the number of elements of the array transducer that are
used at one time by the beamformer, which is also called the channel count. A 64-channel
ultrasound system is capable of better lateral resolution than a 48-channel system.
These two targets
cannot be resolved
separately
Received echo
Transmitted ultrasound pulse
These two targets can
be resolved separately
Received echo
Transmitted ultrasound pulse
Figure 11.18 Schematic showing axial
resolution and its dependence on the
spatial extent of the ultrasound pulse.

11. PHYSICS OF ULTRASOUND IMAGING
79
An important consideration in ultrasound imaging is that the lateral resolution is not
uniform across the image. Locations that are near the transmit focus have better lateral
resolution than locations outside of the focal zone. Therefore, when setting the imaging
parameters, it is important to set the focal zone in the region of interest to get optimal
lateral resolution. Typically, the lateral resolution is worse than the axial resolution in
ultrasound imaging.
Elevational Resolution
The elevational resolution is the minimum distance between two resolvable point tar-
gets in the elevation direction, which is perpendicular to the image plane. The eleva-
tional resolution can also be thought of as the slice thickness of an ultrasound image. In
conventional 2D imaging transducers, there is limited ability to focus in the elevational
plane, and the focusing is primarily achieved using an acoustic lens on the transducer.
Therefore, typically the elevational resolution is worse than the lateral and axial reso-
lution. In recent years, some manufacturers have introduced 2D array transducers for
real-time 3D imaging that have the ability to focus in the elevational plane as well.
Currently, these 2D array transducers are primarily being used for cardiac imaging.
Contrast Resolution, Dynamic Range
Contrast resolution is defi ned as the minimum difference in echogenicity of two differ-
ent tissue types that can be separately resolved. The acoustic physics determines the
fundamental limits on the contrast resolution, but the image postprocessing applied by
the ultrasound system can affect the contrast resolution as well. There are a few imaging
settings that can be adjusted to improve contrast resolution if needed. The fi rst is the
gray scale color map. Different color maps can provide different amounts of contrast
enhancement to different gray levels. Another control available on some ultrasound sys-
tems is dynamic range. It determines the number of gray levels that are represented in
the image displayed on the screen corresponding to a range of echogenicities. Increasing
These two targets
cannot be resolved
separately
These two targets can
be resolved separately
Figure 11.19 Schematic showing lateral resolution and its dependence on the ultrasound beam shape.

ULTRASOUND-GUIDED CHEMODENERVATION PROCEDURES
80
the dynamic range setting enables more brightness levels to be displayed, while decreas-
ing the dynamic range setting limits the number of brightness levels and increases the
contrast between two closely separated brightness levels.
ADVANCED TECHNOLOGIES
In recent years, several new technologies have been developed and introduced into
commercial products. Many of them are relevant to musculoskeletal imaging. Currently,
these technologies are available on premium ultrasound systems, but increasingly some
medium-range systems come equipped with these features. Some of the technologies
most relevant to musculoskeletal imaging will be briefl y described here.
3D Imaging
Conventional ultrasound transducers image a 2D tomographic slice through the body.
This can be extended to perform 3D imaging using two approaches: (a) a mechanical
motor can be used to move the transducer in the elevational plane producing a stack
of 2D images, each acquired at a different elevational location. This stack of images can
then be reconstructed into a 3D volume (Figure 11.20); (b) a transducer consisting of a
2D matrix of elements can be electronically steered in the elevational direction as well.
This enables the ultrasound beam to no longer be confi ned to a 2D plane. There are pros
and cons of each approach. The mechanically-scanned 3D transducers are signifi cantly
less expensive than the 2D matrix transducers. The image quality possible in the 2D
matrix transducer is superior, because of the possibility for elevational focusing. Both
these types of probes can produce real-time 3D (or 4D) imaging. Currently, there are
several disadvantages of 3D imaging that limit its practicality. The visualization and
interpretation of 3D volumes in multiple planes can be quite challenging. 3D imaging
can be time consuming. 4D imaging is slower compared to real-time 2D imaging. High
frequency 3D transducers currently available commercially are mechanically swivelled
and cumbersome, while 2D matrix transducers are expensive. The utility of 3D for mus-
culoskeletal imaging is yet to be fully established, but several opportunities exist:
1. Improved guidance during interventional procedures
2. Ability to measure volume of small muscle groups, or better visualize fl uid accumu-
lation, muscle tears
3. Ability to better visualize anatomical relationships between muscle, tendons, and
bony landmarks
Panoramic Imaging
A limitation of conventional ultrasound imaging is the limited fi eld of view that corre-
sponds to the physical footprint of the transducer. Often, it may be desirable to visualize
an extended fi eld of view to capture a larger muscle. Most ultrasound systems nowa-
days come equipped with the capability to perform panoramic imaging. This technol-
ogy utilizes image-processing techniques to stitch together multiple images acquired

11. PHYSICS OF ULTRASOUND IMAGING
81
Figure 11.20 3D image of the upper
trapezius muscle acquired using a
mechanically-scanned 3D transducer.
Figure 11.21 Panoramic image of the rectus femoris muscle.

ULTRASOUND-GUIDED CHEMODENERVATION PROCEDURES
82
into a panorama. While this technique can enable qualitative visualization of a large
muscle (Figure 11.21), there may be errors in the quantitative estimation of the muscle
volume unless the scan is performed carefully. Panoramic imaging can have interesting
applications in musculoskeletal ultrasound research (4).
Compound Imaging
Array transducers can be electronically steered, and can be programmed to insonate an
object from multiple, slightly different viewpoints (5). These images are combined to
give the spatially compounded image. There are several advantages of this approach in
musculoskeletal applications:
1. It enables improved visualization of longitudinal structures, which are sensitive to
insonation angle. These structures, such as needles, muscle fascicles and so on, can
be visualized much better with compound imaging.
2. Compound imaging can lead to speckle suppression. The speckle pattern changes
when viewed from different insonation angles, thus combining these random varia-
tions in the speckle pattern can lead to suppression in the speckle without affecting
the interfaces being imaged.
Elastography
Elastography is a technology that has been in development in academic research lab-
oratories since the early 1990s but is only now being introduced into commercial
systems. Elastography is designed to assess the mechanical properties of tissue in
response to some perturbation. There is signifi cant interest in quantifying mechanical
properties of skeletal muscle, and elastography has been shown to be a useful tool
toward this goal (6). Based on the type of perturbation, elastography techniques can
be broadly divided into two groups (7): (a) static elastography, where the perturba-
tion is a static compression applied using the ultrasound transducer, or some other
means. In response to this static stress, the underlying tissue deforms. The degree of
deformation, or strain in the tissue, is related to the stiffness of the tissue. The strain
distribution in the tissue in response to the applied compressive stress can be esti-
mated using ultrasound techniques. This method is also called strain imaging; (b)
dynamic elastography, where a dynamic perturbation is applied to the tissue, and
the mechanical response of the tissue is estimated using ultrasound. There are sev-
eral classes of dynamic elastography methods. One class of methods uses an external
vibration and estimates the distribution of vibration amplitudes in the tissue as a
measure of the elastic modulus. These methods are known as sonoelastography. A
second class of methods uses an externally-induced transient perturbation in the tis-
sue and measures the resulting shear wave that propagates through the tissue using
ultrasound. The speed of the shear wave is directly related to the elastic properties
of the material. This class of methods is called transient elastography. A third class
of methods uses the radiation force of the ultrasound to induce a local perturbation
deep inside the tissue, and measures the speed of the resulting shear wave. Muscle

11. PHYSICS OF ULTRASOUND IMAGING
83
tissue is highly anisotropic, and shear waves preferentially propagate along the mus-
cle fi bers. This class of methods is broadly called shear wave imaging. These methods
have been used to quantify the elastic modulus of muscle both in relaxed as well as
contracted state.
REFERENCES
1. Shung KK, Zipparo M: Ultrasonic transducers and arrays. IEEE Eng Med Biol Mag. 1996;15:20–30.
2. Duck FA: Physical properties of tissue: a comprehensive reference book. London, UK: Academic Press; 1990.
3. Barnett SB, Ter Haar GR, Ziskin MC, Rott HD, Duck FA, Maeda K: International recommendations and guidelines for
the safe use of diagnostic ultrasound in medicine. Ultrasound Med Biol. 2000;26:355–366.
4. Stokes OM, Theobald PS, Pugh ND, Nokes LD: Panoramic ultrasound to measure in vivo tendo achilles strain. Foot
Ankle Int. 31:905–909.
5. Entrekin RR, Porter BA, Sillesen HH, Wong AD, Cooperberg PL, Fix CH. Real-time spatial compound imaging: appli-
cation to breast, vascular, and musculoskeletal ultrasound. Semin Ultrasound CT MR. 2001;22:50–64.
6. Gennisson JL, Deffi eux T, Mace E, Montaldo G, Fink M, Tanter M: Viscoelastic and anisotropic mechanical properties
of in vivo muscle tissue assessed by supersonic shear imaging. Ultrasound Med Biol;36:789–801.
7. Greenleaf JF, Fatemi M, Insana M: Selected methods for imaging elastic properties of biological tissues. Annu Rev
Biomed Eng. 2003;5:57–78.

Trunk
Katharine E. Alter and Steven M. Skurow
Since the fi rst reported use of ultrasound (US) for musculoskeletal (MSK) diagnostic
imaging in 1958, US machines have evolved from bulky freestanding machines that
produced grainy, still images of limited resolution to portable cart-based and lap-
top style units capable of producing highly detailed real-time images, which rival
or exceed those obtained by MRI (1). Modern US systems use advanced transducer
technology and sophisticated signal/image processing to create exquisitely detailed
images. Today, US imaging is the imaging tool of choice for a signifi cant proportion
of MSK diagnostic imaging and many procedures (2). Evidence supports that US
guidance is reported to improve the accuracy of a variety of procedures including
MSK injections and muscle/structure targeting for chemodenervation procedures
(3–12).
Advances in US technology such as real-time/live imaging, Doppler modes, com-
pound spatial imaging, extended fi eld of view (FOV), and the exquisitely detailed
images produced by 3D, 4D modes have expanded the clinical/research applications
and imaging possibilities of US. Modern systems are also highly automated with user
friendly control systems that speed data collection and image processing. The addi-
tion of quantitative or semiquantitative signal processing has expanded the role of US
for diagnostic imaging both for clinical and research applications (11). Despite these
advances, US imaging remains highly operator dependent. It is imperative that clini-
cians who utilize US for diagnostic imaging, research, or procedural guidance (includ-
ing chemodenervation) be familiar with the basics of US physics and the details of the
machine and control operations. This knowledge is required to scan, optimize imaging,
and perform procedures (4,13).
This chapter will cover the basics of US machine and control operations
US IMAGING SYSTEMS
All US systems/machines consist of a processor, monitor, and various transducers. The
specifi c features of individual US machines including their size, computer/processor fea-
tures, and the transducer capabilities must all be considered prior to purchasing a unit. A
machine must meet the diverse scanning needs of all the clinicians/specialties who will
be using the machine. The equipment requirements for chemodenervation procedures
falls under the machine/transducer requirements for MSK US. This requires a sophisti-
cated processor and transducers of various sizes and frequencies (13). At minimum, two
transducers are generally required for chemodenervation procedures. One for superfi cial
12
Instrumentation and Knobology

12. INSTRUMENTATION AND KNOBOLOGY
85
structures and the other for deeper muscles or other structures. Specialized transducers
may be required for other chemodenervation applications including gastrointestinal (GI),
genitourinary (GU), or ear, nose, and throat (ENT) procedural guidance.
US IMAGING SYSTEMS
US Unit/Machine
US machines range from large, less portable cart-based systems to small highly portable
laptop-based systems (Figures 12.1–12.3).
The cost of a US machine ranges from $20,000 to over $300,000, depending on the
features and the number of transducers purchased. Modern machines all include a
computer/processor, which controls machine operations, transducer/signal transmis-
sion, pre and postsignal processing and image optimization. In B-Mode imaging, the
onboard machine uses sophisticated technology to process refl ected sound waves cre-
ating 2D black and white real-time images. For most physicians performing US-guided
chemodenervation procedures, the image quality and features provided by a mid-range
Figure 12.1 Cart-based premium Phillips IU 22 US Machine,
Phillips Bothell, WA.

ULTRASOUND-GUIDED CHEMODENERVATION PROCEDURES
86
machine (with at minimum two transducers) is adequate. If, however, the machine is
to be used by other specialists, for other diagnostic/procedural tasks or for research
purposes, a higher end unit with more advanced features and additional transducers
may be required (3,12).
Before purchasing a machine, physicians must personally use a machine during
a live demonstration, preferably in a side to side comparison with other machines
that are being considered. The proprietary differences between manufacturers and
machines lead to signifi cant differences in features, quality of images produced, and
ease of use. These differences mandate the demonstration/evaluation of equipment
(13,14).
Monitors/Screen Size
Older US systems used freestanding CRT monitors, which were large, bulky, and
reduced portability of US systems (15). In modern machines including cart-based
or laptop style systems, the screen is integrated into the system. Either a computer
screen or fl at screen monitor is provided on which US images are displayed/viewed.
When purchasing a machine, screen size is an important consideration. Compact
highly portable units necessitate a smaller screen size than that typically provided
with large cart-based systems. The screen size of the unit must provide adequate
visualization of the structures to be imaged; in general, most clinicians prefer a
Figure 12.2 Esaote MyLab US unit, Esaote.

12. INSTRUMENTATION AND KNOBOLOGY
87
larger screen. Having a machine demonstrated, in the setting/environment where it
will be used is useful.
TRANSDUCERS
The importance of selecting the correct transducer for diagnostic US scanning or
US-guided procedures (including chemodenervation) cannot be overstated. Transducers
come in a variety of styles, shapes, and frequencies and are selected based on scanning
requirements (see Figures 12.4–12.7) (12,13,16). Transducers (piezoelectric crystals) con-
vert electrical pulses to mechanical/sound waves, which are then transmitted to the
body. The same transducer, (i.e., piezoelectric crystals) converts the returning mechan-
ical wave/vibrations to an electrical potential/signal. This signal is conveyed to the
computer to be processed and displayed. The frequency of the transducer determines
depth of penetration and resolution (Table 12.1). During scanning and procedures clini-
cians should use the highest frequency transducer that provides adequate depth of pen-
etration to visualize a target (17).
The “foot-print” of the transducer refers to the area of the transducer that comes in
contact with the patient. The US beam width, however, is considerably narrower than the
short axis of the footprint. Footprint also determines the width of the US beam. (Figures
12.7b and 12.8a–c) The displayed image represents only a thin slice or cut through a
structure, similar to that obtained with computed tomography (CT). To compensate for
the narrow width of the beam, the operator must scan, moving the transducer up and
down along the length and back and forth across the width of a structure to visualize
structures in their entirety (13,16,17).
Transducer Resolution
To discriminate between two structures requires adequate spatial resolution. In general,
the higher the frequency, the better the spatial resolution. However, frequency alone
(b)
(a)
Figure 12.3 (a) Terason 3200, split screen, torn calf muscle (b) Cart/Mobility System for Terason 3200
Machine.

ULTRASOUND-GUIDED CHEMODENERVATION PROCEDURES
88
does not determine spatial resolution. The details of sound generation, detection, and
signal processing all infl uence the spatial resolution achieved by individual machines/
transducers. Because of this, not all transducers or machines are created equal, meaning
that transducers of the same frequency, from different manufacturers, may provide very
different images. This fact makes it vital to test any machines/transducers during a live
demonstration prior to purchase. This should be done preferably in a side to side com-
parison with several machines (13,18).
Figure 12.6 Pediatric abdomen transducer.
(a)
(b)
(c)

Figure 12.4 Linear transducers (a) 12–5, (b) 17–5, (c) Hockey Stick.
Figure 12.5 Curvilinear transducer.

12. INSTRUMENTATION AND KNOBOLOGY
89
Axial Resolution
Axial resolution is the ability to discriminate between two points/structures, one lying
above the other along the depth axis (Y-axis). Axial resolution is proportional to trans-
ducer frequency.
Lateral Resolution
Lateral resolution is the ability to distinguish between two adjacent structures
lying side by side, perpendicular to the sound beam (X-axis). As with axial resolu-
tion, lateral resolution improves with higher frequencies. Lateral resolution is far
more dependent on focal zone placement than is axial resolution, however, apples
to apples, the best lateral resolution occurs within a focal zone (see, Focal Zone,
below).
Figure 12.7 (a) Linear and cardiac specialty transducers (b) 3D transducer.
Table 12.1 Transducer frequency, depth of penetration
M Hz
Depth/Penetration
Application
3
12 to 20 cm
OB/GYN
5
12 to 15 cm
Deep muscles: Piriformis, psoas, iliacus, quadratus lumborum
7.5
8 to 10 cm
Moderately deep: Thigh, leg
10
5 cm
Moderately deep: Calf, arm, pediatric (leg/thigh)
12
3.5 cm
Superfi cial: Calf, forearm
17
2 cm
Superfi cial: Face, intrinsic hand/foot muscles