Monthly Archives: June 2014

Continuing Education

Expositions. The environment in which clinical engineering functions changes daily as new technologies such as telemedi­cine, robotics, and wireless local area networks (LAN) are in­troduced into the clinical setting. In this dynamic field contin­uing education is the rule. One method of obtaining this education is by attending technical expositions and profes­sional organization meetings. The Association for the Ad­vancement of Medical Instrumentation (AAMI) and the Amer­ican Society of Healthcare Engineering (ASHE) hold meetings and expositions that expose attendees to the latest medical instrumentation being introduced into the marketplace. The pulse of the health-care industry can be sampled in a rela­tively short time by attending roundtable discussions and member and industry presentations. Courses are provided in regulatory requirements, medical devices, instrumentation repair and maintenance, clinical engineering and BMET pro­fessional certification preparation, and clinical engineering management. Technical information is also presented at monthly meetings of the Institute of Electrical and Electron­ics Engineers (IEEE) Engineering in Medicine and Biology Society, as well as local clinical engineering and BMET orga­nization meetings.

Service Training. As new medical equipment is acquired, employee technical knowledge must be updated with regard to its operation, preventive maintenance, and servicing. Training is available from the manufacturer or from indepen­dent schools. Training can sometimes be included in purchase requisitions and request for quotations (RFQs) for new equip­ment. Service training not only benefits the clinical engineer and BMET involved in maintaining this equipment, it also benefits the equipment user each time clinical engineering is called on to assist them with equipment-related questions. Formal service training can be expensive. In addition to tu­ition there are travel and lodging expenses. To supplement, but not replace service training, clinical engineering staff can attend equipment operator training provided by vendors for clinical users of medical equipment, within their own institu­tion. Training can also be obtained using the expertise avail­able within the clinical engineering department (Fig. 1).

Self-Study. Formal training can be supplemented with self­study of technical journals, periodicals, and trade publica­tions, as well as equipment operator and service manuals, VCR training tapes and computer-based training programs. A clinical engineering library provides an invaluable tool for the clinical engineering staff and for other health-care work­ers (physicians, nurses, laboratory technicians) to whom clini­cal engineering services are provided. Libraries could include technical video, equipment operator and service manuals, and technical publications (books and magazines). Such material also allows staff to keep pace with changes in regulatory re­quirements and biomedical standards.

Safety Training. Employee right-to-know and safety train­ing that discusses the hazards encountered in the workplace is also necessary. This includes subject matter related to blood-borne pathogens, hazardous materials, proper protec­tion when entering patient-care areas (gloves, masks, etc.), environmental hazards, fire hazards, patient’s bill of rights, and other items. The latest trend makes use of interactive computer program modules. This allows training at a time convenient to the employee and no longer requires attendance at lengthy seminars.

Certification

Certification is provided for clinical engineers [Certified Clini­cal Engineer (CCE)] and biomedical equipment technicians [Certified Biomedical Equipment Technician (CBET)] by ex­amining boards guided by the International Certification Commission for Clinical Engineering and Biomedical Tech­nology. The Commission is composed of health-care commu­nity members including engineering, medical, industrial, and governmental groups and agencies. Certification provides for­mal recognition that an individual has mastered a body of knowledge that is useful in job performance. This knowledge, which is both theoretical and practical, includes theory of op­eration of medical equipment, physiological principles, and safety issues related to medical equipment (8).

Clinical engineering certification requires passing a writ­ten exam (multiple-choice and essay questions), and an oral interview, aimed at determining the candidate’s depth and breadth of experience. BMET certification requires passing a written multiple choice examination. The Association for the

Figure 1. In-service education, ventilator tester. Inservice education can be provided by manufacturers and vendors, as well as by clinical engineering staff. Here a clinical engi­neering supervisor is providing training to other clinical engineers in the use of an auto­mated ventilator tester. Such devices are used during preventive maintenance and repair. They reduce the number of individual test in­struments required as they integrate several test functions into one device. This reduces service time leading to more rapid equipment turnaround. Such education also helps to sat­isfy JCAHO training requirements for clinical engineering staff.

Advancement of Medical Instrumentation (AAMI) assists can­didates by providing certification training courses and study materials.

Certification renewal requires demonstration of continued training. Points are assigned and accumulated for various ac­tivities that contribute to one’s ability to do his job.

CLINICAL ENGINEERING

OVERVIEW Definition

Clinical engineering is a relatively new profession. The term clinical engineering was coined by Cesar A. Caceres, M. D. in 1967 to describe a George Washington University Medical School program in which he envisioned engineers and physi­cians working together to provide better patient care. This program, which was not medicine, engineering, nor statistics, contained elements of all of these disciplines. As the program focus was to be patient oriented, he chose to couple the term clinical with engineering to describe it, so as to distinguish it from the research-oriented activities of biomedical engi­neering (1).

In 1992 the American College of Clinical Engineers (ACCE) defined a clinical engineer as ‘‘a professional who sup­ports and advances patient care by applying engineering and management skills to healthcare technology” (2). The Clinical Engineering Board of Examiners of the International Certifi­cation Commission for Clinical Engineering and Biomedical Technology endorses this definition.

Environment of Patient Care

At the heart of clinical engineering is the concept of providing engineering expertise to ensure that the environment of pa­tient care (EC) is safe for both patient and clinician. Medical equipment used for patient care comprises a large part of this environment. It must be safe, efficacious (performing the function for which it was intended), and cost effective. Clini­cal engineers are the professionals who provide technical sup­port services to ensure this. Their knowledge is invaluable to health-care provider institutions, such as hospitals, nursing homes, clinics, medical and dental offices, and ambulatory care centers.

Clinical engineers are trained in engineering principles, basic sciences, the life sciences and patient-care principles. They are knowledgeable of regulatory agency requirements, health-care codes, and medical equipment standards. This di­verse body of knowledge and experience enables clinical engi­neers to understand how technology and patient care interact within the clinical setting (3). This knowledge is vital when medical equipment is integrated into the environment of pa­tient care. It allows clinical engineers to interface medical systems to the patient, and to other medical systems includ­ing those used for data collection (4).

Employment

Primarily employed by health-care provider institutions (as part of in-house or shared-service clinical engineering depart­ments), clinical engineers are also employed by original equipment manufacturers (OEM), third-party independent service organizations (ISO), independent testing laboratories, clinical engineering consulting firms, regulatory agencies, technical publishing houses, law firms, and academic institu­tions.

Clinical engineers hold positions in management, engi­neering, medical equipment sales, equipment test and evalua­tion, field service, health-care regulation, technical publish­ing, and education. They serve as expert witness of patient incidents and provide input to legislative bodies. They also serve on curriculum advisory committees on which clinical en­gineers from both industry – and hospital-based clinical engi­neering programs sit, allowing curriculum to keep pace with the most recent industry trends.

Biomedical Equipment Technicians

Closely associated with clinical engineers are biomedical equipment technicians (BMET). BMETs are skilled techni­cians who are specially trained to work with medical instru­mentation. Although BMETs focus their activities on the re­pair and maintenance of medical equipment, they are called on to inspect, install, and modify medical devices, as well as to provide guidance in proper equipment usage and safety. Some take on managerial responsibility, supervising other BMETs.

Education

Clinical Engineering. As a minimum, new practitioners in clinical engineering require a bachelor of science degree in engineering. This degree should be obtained from an institu­tion that is accredited by the Accreditation Board for Engi­neering Technology (ABET).

Formal clinical engineering curricula are offered by col­leges and universities. Bachelor’s, master’s, and doctoral de­gree programs are available in biomedical and clinical engi­neering. The important difference between the curriculum for these programs and the curriculum for other engineering dis­ciplines is the mix of engineering and life sciences that it of­fers. Included are traditional engineering courses (electrical, mechanical, chemical, computer engineering), as well as courses in the physical sciences, life sciences (biochemistry, biology, physiology, and anatomy), mathematics, humanities, and management (5).

It is also possible to enter the clinical engineering field with a traditional bachelor of engineering degree (such as electrical engineering) and then acquire the necessary life sci­ences knowledge by taking supplementary courses, on-the-job training, and self-study.

Clinical engineers are qualified to pursue advanced de­grees in such diverse fields as medicine, law, business admin­istration, health-care management, and technology as­sessment.

It should be noted that prior to formal degree programs becoming available in the 1970s, early leaders in the field en­tered with backgrounds in the physical sciences or life sci­ences and are considered to be “grandfathered” into the pro­fession.

The following sources of information are useful (6) :

Directory of Engineering and Engineering Technology: Un­dergraduate Programs from the American Society for Engineering Education

Peterson’s Guide to Undergraduate Programs in Engi­neering and Applied Sciences

Peterson’s Guide to Graduate Programs in Engineering and Applied Sciences

Biomedical Equipment Technician. BMET education leading to an Associate in Applied Science Degree (AAS) in Biomedi­cal Engineering Technology is offered via seven accredited programs in the country. One such two-year program is of­fered by the State University of New York, College at Farm – ingdale. This program provides balanced course work in elec­tricity and electronics, chemistry, physics, physiology, and biomedical engineering technology. Students have the option of continuing their education an additional two years earning a Bachelor Degree in Electrical Engineering Technology (7).

Coagulative Ablation Therapy

The development of coagulative ablation therapy over the past decade has revolutionalized the practice of car­diology, gynecology, and urology. For many of the diseases, surgical intervention has been the principal method of treatment, although alternatives to surgery have been sought in an effort to reduce the cost and morbidity of treatment (107,108,109,110,111,112,113,114). Minimally invasive catheter ablation offers several potential benefits: Long incisions are replaced with a puncture wound, major cardiac and pulmonary complications from general anesthesia are side-stepped, and the need for postoperative intensive care is significantly reduced and, in many cases, offers a complete or lasting cure. It also has important advantages over drugs that are merely palliative. It avoids the side effects, expense, and inconvenience of chronic drug therapy, often with only partial success.

Energy sources that can be used for ablation therapy include RF, microwave, laser, and ultrasound. Before describing the RF and microwave technology, a brief discussion of laser interactions with tissue is given since it is used for the thermal therapy.

The effect of lasers on tissue is related to fluence (the product of power density and duration of irradi­ation). Thermal ablation involves the delivery of laser power for periods lasting tens to hundreds of seconds. The temperature can easily reach 100°C, depending on the delivery system. The desire to limit excessive tissue damage (myocardium, uterine wall, etc.) has led to several innovative technologies for laser thermal interven­tion. However, the complication of perforation or dissection is a cause for special attention in laser sources (115,116,117,118,119,120,121,122,123). Nevertheless, laser catheter irradiation (e. g., Nd:YAG, 1064 nm, 15 W to 50 W, spot diameter 2.0 mm to 2.5 mm) has been shown to produce lesions selectively in the targeted segment of the right ventricular conduction system in dogs, and the method can be performed in a controllable manner (120). Endometrial ablation techniques using laser coagulation under direct hysteroscopic control have been at­tempted with varying success (115,116,117,118,119). The variability arises principally from the unpredictable nature of induced thermal injury and perforation of the uterine wall. Laser coagulation prostatectomy is used to improve urinary flow rates (121,122,123). While results are comparable to standard electrocautery resection, the procedure can be enhanced by modifying the laser regimen and the spatial distribution of lesions. Since medical applications of laser are discussed elsewhere, it will not be addressed further in this section.

RF and Microwave Ablation. In RF thermal ablation therapy, the current flows between a small electrode inside the body to a large grounded electrode on the surface. The current rapidly diverges from the small electrode, so that current density is the highest at the electrode-tissue interface. The tissue’s resistance to current flow results in termal lesions: desiccation and coagulation of tissue in direct contact with the electrode. The desiccated and coagulated tissue would raise the resistance to current flow, impede effective tissue heating, and limit the size of RF-induced lesions. Lesion beyond the immediate vicinity of the electrode-tissue interface occurs as a result of passive heat transfer from the thin high-temperature region. Investigations have shown that RF-induced lesions increase rapidly in size during the initial period of power application; then the rate of increase diminishes rapidly as the resistance at the electrode-tissue interface rises and the current flow falls (124,125,126,127). For this and reasons described below, studies comparing the power deposition patterns of RF and microwave catheters have shown that the absorbed microwave energy could be 10 times higher than RF at the same tissue depth (128,129).

The frequencies of most interest to microwave ablation are 915 MHz and 2450 MHz. Typical values of microwave dielectric permittivity and conductivity at 37°C are given in Table 2. The biological tissues of interest to ablation therapy can be classified into three major groups according to their water content. The group with very high water content includes blood, uterine lining, or physiological fluids. The second group is of moderately high water content and includes muscle or cardiac wall. The third group is made up of tissues with low water content such as bone, fat, or desiccated tissue. It can be seen that there is a modest change in dielectric constant and conductivity as a function of frequency. However, differences among the tissues are

Table Й. Dielectric Constant риД Conductivity of Blood, Muscle, ят*Н Fat Tissues at 37C for 915 and 3460 МНй-Твд Frequencies 0/ Most Interest to Coagulative Ablation ТЬегіаруй

D ielectric с [instant

Conductivity (й/ш)

Frequency (MHz)

Blood Muscle

Fat

Blood

Muscle

Fat

915

60 61

5.6

1.4

1.6

0.10

2450

58 49

5.5

2.1

2.2

0.16

5 From Ryf 36.

quite large. The higher 2450 MHz frequency is chosen because at this frequency the dielectric constant for blood is 20% higher than that for muscle, and the dielectric constant muscle is about 800% higher than that for fat. While conductivities of blood and muscle are approximately the same, they are about 300% higher than that of fat. As the microwave radiates into the tissue medium, energy is absorbed and converted to heat by dielectric loss. This absorption will result in a progressive reduction of the microwave power intensity as it advances in the tissue. The time rate of heating and spatial distribution of radiated microwave energy at any given moment in time are direct functions of SAR and antenna radiation pattern, respectively.

The reduction is quantified by the depth of penetration; a measure of the distance through which the intensity of a plane wave field is reduced to 13.5% of its initial level in a medium. At 2450 MHz, the depths of plane wave penetration for blood, muscle, and fat are 19 mm, 17 mm, and 81 mm, respectively (35,36,37). For microwave catheter antennas which do not have plane wavefronts, the penetration depth is reduced according to the specific antenna design. Nevertheless, these values clearly suggest that microwaves can deposit energy directly into distant tissues. Furthermore, the difference in the dielectric permittivity yields a depth of penetration for tissues with low water content about four times deeper for muscle or higher water content tissue at 2450 MHz. This means that a microwave field can propagate more readily through and be absorbed less by low water content tissues than that of high water content. It also implies that microwaves can propagate through intervening desiccated tissue or fat to deposit energy directly into more deeply-seated tissue.

Cardiac Ablation for Tachyarrhythmia.

For a significant portion of patients suffering from tach­yarrhythmias, available drug therapy has been found unsatisfactory because of a lack of meaningful response or unacceptable side effects (107,108,109). In some cases, these patients can be meanaged by open-heart surgery. Percutaneous catheter ablation of arrhythmogenic foci inside the heart is a potentially curative mode of treatment. Indeed, RF ablation has emerged as an effective therapy for many supraventricular tachycardias and has become accepted as the standard treatment for arrhythmias associated with the Wolf-Parkinson – White syndrome (107,108,130). Typically, the catheter is inserted percutaneously into the femoral vein and, under the guidance of a fluoroscope, is then advanced to inside the heart chamber. The cardiac conducting tissue responsible for the tachycardia is identified with the aid of endocardiac electrograms. A burst of RF energy is delivered through the electrodes to thermally ablate the cardiac conducting tissue responsible for the tachycardia and restores the heart to its normal rhythm.

Rapid and reliable mapping of the endocardiac electrogram for identification remains a technical chal­lenge. Also, the lesions induced by RF current is quite small and shallow (125,127,131). Increasing the output power to heat tissue at a distance often results in excessive temperatures at the electrode-tissue interface without the desired enlargement of lesion size (126,132). Note that temperature-guided RF catheter ablation with very large distal electrodes can be used to improve lesion size (133). The impedance of the ablating elec­trode would rise due to poor coupling between the electrodes and adjacent tissue; desiccated and coagulated tissue raises the resistance to current flow, thwarts effective tissue heating, and limits the size of RF-induced lesions.

There is a need for energy sources that can produce larger and deeper lesions than RF currents. Large lesions are required for certain types of cardiac ablation to cure ventricular tachycardias secondary to coronary artery disease, for example, and arrhythmias due to reentry located deep in the myocardium in particular. The radiating and dielectric heating features of microwave energy theoretically may be useful for ventricu­lar ablation. The interaction of microwaves as mentioned earlier can result in a greater volume distribution of energy and deeper penetration. The feasibility of ablating the atrioventricular (AV) junction in dogs with microwave catheter antennas has been shown both in vitro (134,135,136) and in vivo (137,138,139,140). Fur­thermore, using fresh bovine hearts and closed-chest dogs, the feasibility of a larger (4 mm long) split-tip catheter antenna has been demonstrated for ablation treatment of ventricular tachycardia (141,142). The re­sults suggest that if the lesion size is sufficiently large, it would be possible to ablate a ventricular tachycardia focus using this split-tip microwave catheter antenna system. In addition to the split-tip catheter antenna, microwave antennas reported for cardiac ablation include monopole, helical coil, cap-slot, and cap-choke de­signs (134,135,136,137,138,139,140,141,142,143,144,145,146). A drawback of some catheter antennas is that a considerable amount of microwave energy is reflected by the antennas to the skin surface and is deposited at the point of antenna insertion into the blood vessel. The problem has been addressed by integrating a sleeve or choke in the antenna design (134,135,136,139,141,142,143,144).

It is noted that the feasibility of using ultrasound for cardiac ablation was investigated and a catheter mounted transducer has been reported (147,148).

Endometrial Ablation. Hysterectomy is performed to surgically remove the uterus in order to stop intractable bleeding or menorrhagia (149,150). Endometrial ablation is a relatively new treatment for menor­rhagia and is a reliable alternative treatment for patients with dysfunctional uterine bleeding (151,152,153). It is superior to hysterectomy in terms of operative complication and postoperative recovery. While still in the beginning stages, RF and microwave thermal ablation of the endometrium have been reported as efficacious procedures for treatment of abnormal uterine bleeding (154,155,156,157). The technique is easier and quicker to perform than current alternatives. Quantitative measures and patients’ subjective responses suggest that a meaningful fraction of patients treated with RF and microwave ablation experience significant flow reduction. Investigations with microwave energy indicate that a treatment temperature of 55°C is related to signifi­cant reduction or complete elimination of menstrual flow (156,157). Besides the difference between microwave and RF approaches mentioned already, the use of high-intensity RF power (500 W) could produce burns at points where electrocardiographic (ECG) electrodes come in contact with the body (158). Considerably more investigation is needed before microwave or RF ablation can become a safe and efficacious clinical modality.

Treatment of Benign Prostate Hyperplasia. Benign prostatic hyperplasia or hypertrophy is a major cause of morbidity in the adult male. At present, open surgery and transurethral resection of the prostate are the gold standards for treatment of benign prostatic hypertrophy. They can provide immediate relief of obstructive symptoms that remain fairly to extremely durable (159). A new, less invasive procedure uses thermal energy delivered by microwaves (160-165). An early report of a thermal microwave technique from 1985 employed a transurethral microwave applicator. It showed coagulation of the prostate in mongrel dogs and some salutary effects in an initial six patients treated with this device (160). An ensuing study used 2450 MHz microwave energy to treat 35 patients and compared transurethral resection alone to preliminary microwave coagulation followed by transurethral resection of the gland (161). Significant reduction in blood loss by initial treatment with microwave thermal therapy was observed. Numerous reports have appeared since that time on various aspects of both transrectal and transurethral microwave therapy of the prostate using 915 MHz and 2450 MHz energy (162,163,164,165,166,167).

Most of the research in human subjects to date has focused on methods of delivery. Initial attempts to deliver the energy transrectally have not been effective, and injury to the rectal mucosa has occurred due to the difficulty of interface cooling of this organ (166,167). Recent investigations have focused on transurethral delivery of the energy with cooling systems within the catheter to ensure urethral preservation (143,144, 162,163,164,165,168,169). Sensors placed in the microwave antenna maintain temperature on the urethral surface between 43° C and 45° C. It is noted that while the number of treatment sessions and the temperature attained are extremely important predictors of response, sufficient hyperthermia volume is crucial for enhanced efficacy. Virtually no data clearly demonstrating reduction in prostate volume in human subjects have been reported, although most investigators have shown improvement in measured urinary flow rates compared to preoperative studies. Randomized studies comparing microwave thermotherapy to transurethral resection conclude that microwave hyperthermia treatment had a definite therapeutic effect on symptomatic prostatic hypertrophy (169,170,171,172). Thus, microwave thermal ablation of prostatic tissue and enlargement of the urethra with minimal clinical complications offers a therapeutic alternative to surgery in select patients with benign prostatic hyperplasia.

Hyperthermia Treatment for Cancer

Hyperthermia cancer therapy is a treatment procedure in which tumor temperatures are elevated to the range of 40° C to 45° C. The rationale for hyperthermia therapy is related to the ability of elevated temperature to selectively destroy malignant cells. While tumor cells exhibit inherent hyperthermic sensitivity, their response is characterized by the acidic, hypoxic, and nutritionally deprived environment often found in the interior of various tumors (4,5,6). Poor blood perfusion in the interior of a tumor also facilitates heat buildup. Moreover, the cytotoxic effects of some autitumor drugs are enhanced and the cell-killing ability of ionizing radiation is potentiated by hyperthermia serving as a sensitizing agent. Hyperthermia also increases blood-brain barrier permeability (7). The synergism of hyperthermia and ionizing radiation is particularly poignant since it is accomplished by thermal killing of hypoxic cells and cells in S phase (DNA synthesis), which are resistant to ionizing radiation.

Clinical and laboratory results from various countries have indicated a promising future for hyperthermia (8). Its efficacy depends on the induction of sufficient temperature rise throughout the tumor volume. A recent assessment of superficial breast cancer has indicated that local complete response with hyperthermia and ionizing radiation is about 60% compared to 40% with irradiation alone (9). Currently, hyperthermia is still an experimental treatment in the United States for late-stage patients with advance tumors, but it has gained some acceptance in Europe and Japan (10,11,12).

While beyond the scope of this article, whole-body hyperthermia has been employed in some cases to enhance the effectiveness of chemotherapy for patients with systemic metastatic cancer. A variety of conductive and convective heating techniques such as warm air, water, and wax are used (13,14,15). Mild whole-body hyperthermia at 40°C for as long as 10 hours in rats has shown promising therapeutic potentials on primary tumor. Temperature up to 41.8°C is found to be safe and is well-tolerated by human patients for up to 60 min (16,17). While clinical results still remain guarded, they provide a foundation for further exploration.

Monitoring and control of tumor temperature in real time during hyperthermia treatment is essential for effective therapy. While progress in temperature sensing in vivo has been dramatic, considerable advance is needed prior to widespread clinically application of hyperthermia for cancer. Among approaches that may impact this outcome include invasive multipoint sensing (18) and noninvasive magnetic resonance imaging and diagnostic ultrasound temperature mapping (19,20,21).

A prominent problem in hyperthermia treatment for cancer is the generation of heat and control of tem­peratures in tumors. Ultrasonic and electromagnetic energies are the commonly used sources for regional and local hyperthermia. A large number of external and implanted antennas and applicators have been designed to produce therapeutic heating of localized tumors of different volumes in a variety of anatomical sites. Clearly, each modality has its own liabilities. Recently, there was a comparison of temperature distributions obtained in the same tumors, and there was also a comparison of acute and subacute toxicities in patients that were treated with both external ultrasound and electromagnetic applicators (22). It was concluded that there is no preferred modality. The type of applicator should be selected on the basis of specific site and type of the tumor.

Ultrasonic Heating. Absorption of ultrasonic pressure wave in biological tissue is determined by ul­trasound frequency, velocity, and tissue density. Several frequencies between 0.5 MHz and 3.5 MHz have been used for hyperthermia cancer treatment. At these frequencies, ultrasound can penetrate deep inside the body while maintaining the ability to focus energy into the tumor to raise the temperature tumor volume to above the minimum therapeutic temperature. Its clinical application is constrained by the available ultrasound window between the transducer and target tumor, as well as by the presence of bone and soft tissue interfaces in the propagation pathway. Differences in tissue density can give rise to excessive temperature elevation resulting from accumulation of reflected power at these interfaces. For a given set of anatomic and physiologic param­eters, temperature distribution in a tumor is determined by transducer design, scanning pattern, scanning speed, and output power.

Ultrasonic modalities for noninvasive hyperthermia cancer treatment include spot focus transducers and phased arrays. By mechanically scanning a focused transducer around a treatment volume, uniform tempera­ture distribution with a sharp falloff outside the treatment volume may be obtained (23,24,25). However, for large, deep-seated tumors, scanning transducers often produce hot spots proximal to the tumor along the cen­tral axis ahead of the focal plane (26). It should be mentioned that frequency sweeping and transducers with a nonvibrating center can be used to reduce the central hot spot (27). There are two classes of phased arrays that do not require physical movement of the transducer elements (28,29,30,31,32). The class of phased arrays with geometric focusing and spot scanning has features and limitations similar to those of mechanically scanned spot focus transducers (28). In this case, the transducer array is fixed in position and electrical spot scanning is accomplished through adjustment of array element phases which maximize constructive interference at each focal plane.

Alternative array element configuration and phase excitation can avoid hot spots that result from con­structive interference along the array’s central axis (29,30,31,32). Several phased array configurations have been proposed. They include the concentric ring, sector vortex, spherical section, and the square arrays. With proper selection of array element phases, these phased arrays can be operated to directly synthesize, with­out scanning, ultrasonic power deposition patterns for improved localization of heating within the tumor volume. Phased arrays offer another advantage over single focused transducer: They enable electronically programmable treatment planning (33). Pretreatment analysis can provide strategies aimed at satisfying therapeutic requirements for individual patients and specific tumor sites and spare other sensitive anatomic structures. It is noted that recently ultrasonic applicators have been tested for interstitial hyperthermia (34).

Electromagnetic Heating. Various frequencies of electromagnetic energy within the range of 0.05 MHz to 2450 MHz have been used for hyperthermia treatment of cancer. The interaction of electromagnetic fields and waves in biological tissue is governed by (1) source frequency and intensity, (2) antenna or applicator design and polarization, (3) tissue structure, and (4) dielectric permittivity (35,36,37). In thermal therapeutic applications, the final temperature is affected also by tissue blood flow and heat conduction. However, the time rate of heating and spatial distribution of electromagnetic energy at any given moment in time are direct functions of specific absorption rate (SAR or power deposition), which are functions of antenna or applicator design.

At frequencies below a few hundred megahertz [i. e., radio frequencies (RF)], wavelengths in tissue are 100 cm or longer (see Table 1). Power deposition for local tissue heating is characterized by quasistatic displace­ment or conduction currents, and heating comes about through tissue resistance to current flow. At microwave frequencies, wavelengths are much shorter, radiated power dominates, and dielectric loss gives rise to heat production. Power coupling from air into tissue is substantial and can exceed 50%. In addition, the effective depth of penetration can provide useful insight into the performance of various applicators. For example, be­cause of both the focusing ability and the depth of energy penetration, single-contact applicators operating at 915 MHz or 2450 MHz have been used to heat well-localized superficial tumors extending to a depth up to 3 cm to 6 cm.

RF Heating. For noninvasive subcutaneous tissue heating by RF energy, simple capacitive plates and inductive coil applicators have been used. Tissues are positioned between the plates and are heated by dis­placement currents (38,39). A water bolus is often placed between the plate and skin to prevent superficial burns from large electric field concentration near the edges. A limitation of the capacitive applicator is that the

Table 1. Propagation Characteristics of RF ят*Н Microwave Fields anrf Waves in Planar Model of Biological Tissues of High – and Low-Water* Content Biological Tissues at 3742“

Frequency {MHz} ‘

Dielectric

permittivity

Conductivity

(S/rn) ‘

Wavelength in tissue (cm)

Effective depth of penetratiou {cm)

Power coupling coefficieut from air

High

Loav

High

Low

High

Low

High

Low

High

Low

27

ш

20.0

0,61

0.03

241

14.3

77

0.14

0-66

40

97

14.6

0,69

0.03

51

187

11.2

56.6

0.17

0.62

1O0

72

7.5

0-05

27

106

67

34.4

0-22

0-74

433

53

5.6

1.43

o. oe

3.5

26.2

3.6

16.3

0.36

0.Й2

91E

51

5.6

1.60

0.10

4.4

13.7

25

12.6

0.40

О. ЄЗ

2450

47

5.5

2,21

0.16

l. S

5.2

1.7

8.1

0.43

0.94

aFrotn Ref. 35.

electric field is predominately normal to the interface between fat and other tissues. Overheating by as much as 20 times that in muscle can occur in subcutaneous fat greater than 2 cm in thickness. Note that it is possible to treat tumors of patients with subcutaneous fat as thick as 3 cm by precooling the fat prior to the initiation of heating (40).

The common inductive applicator consisting of a planar or “pancake” coil with a small number of turns when placed parallel to the body surface can avoid the excessive heating problem in fatty tissue. Since the induced electric fields form eddy currents that flow parallel to the tissue interface, heating is highest in muscle instead of fat. The heating pattern is toroidal with a null along the axis of the applicator (41). Some recent designs have SARs that do not include a null in the center and are considerably more uniform than that of the planar coil (38,42,43). While inductive applicators are used predominately for superficial treatment, some of the newer applicators can produce effective heating up to a depth of nearly 7 cm.

Several RF applicators have been invented to provide noninvasive heating of deep-seated tumors. These include the large capacitive applicator mentioned previously (44,45) as well as the ridged waveguide (46), helical coils (47,48), and multielement arrays (49,50,51,52). The helical coil applicator is simple in construction, and it provides SAR patterns that vary slowly with radial distance. However, the region to be heated must be located near the center since the axially directed electric field has a maximum near the center of the coil structure. Its performance may be improved by judicious selection of the diameter-to-length ratio and incorporation of external tuning. The multielement array concept has gained considerable utility in the clinic. A primary advantage of multielement array systems is the ability to steer the heating pattern electronically by varying the amplitude and phase of each element, thereby allowing phased arrays operating at RF to be used for selective heating of deep-seated tumors in a variety of anatomic sites.

In particular, the annular phased array system is utilized to heat large anatomical regions such as the thorax, abdomen, and pelvic area (49,50,53). Regional heating is frequently complicated by systemic hyperthermia and hemodynamic compensation and by excessive heating or adjacent normal tissue struc­tures, especially the bone-tissue interface. However, recent advances using feedback algorithms and adaptive software modifications to control the amplitude and phase of each element showed that it is possible to max­imize the SAR at a target tumor position in a complex anatomy and simultaneously minimize or reduce the power deposition at locations where undesirable hot spots may occur.

A novel noninvasive or minimally invasive concept using ferro – or paramagnetic compounds for intracel­lular hyperthermia treatment of both primary and metastatic cancers was first proposed in the late 1950s. Earlier studies have demonstrated both the preferential accumulation of submicron-sized magnetic particles (magnetites) in tumors and the feasibility of selective heating using 0.24 MHz to 80 MHz RF magnetic fields. Current investigations (54,55,56) are directed toward cellular uptake of fluidized magnetic particles, bounding of magnetite with targeting activity towards cancer cells, and the hyperthermic effects of fine magnetic parti­cles on tumor cells in vitro. It is expected that a magnetite-labelled antibody may soon be available clinically as a therapeutic agent for hyperthermia treatment of cancer.

A related technique for RF hyperthermia involves implanted ferromagnetic seeds activated by externally applied 0.05 to 2 MHz magnetic fields. Heating is produced by eddy currents induced on the surface of the implant and is therefore dependent on the permeability of the thermoseed material (57,58,59,60,61,62). Using Curie temperatures close to the maximum temperature desired in the tissue, the ferromagnetic seeds can be designed to provide thermal self-regulation so that a constant tumor temperature can be maintained throughout the treatment regime. Since volume tissue heating is by passive thermal conduction, these 0.1 mm to 1.0 mm diameter thermoseed of various length must be implanted closely. Nevertheless, under certain conditions this invasive seed implant method like the interstitial RF electrodes and microwave antennas to be discussed later may be preferable for local hyperthermia of deep-seated tumors. It is noteworthy that ferromagnetic seed hyperthermia in combination with other modalities are used in the control of ocular tumors in animals (63,64). Recently, multifilament seeds such as the palladium-nickel (PdNi) thermoseeds have gained interest because of a more effective power deposition than solid seeds (65).

For some deep-seated tumors or tumors of large volume, interstitial techniques have been employed to generate the desired hyperthermic field. RF electrodes operate in the frequency range of 0.5 MHz to 1 MHz (66,67,68,69). The advantages of interstitial techniques are safe (without skin burn) and more uniform heat distribution within the tumor. RF current flowing between pairs of needle-like bare electrodes is dissipated by the ohmic resistance of tissue and is converted to heat. The temperature distribution produced is strongly dependent upon blood flow in the tissue and spacing between electrodes. Most clinical applications require an array of these electrodes spaced at 1.0 cm to 1.5 cm intervals in parallel for optimal temperature uniformity. Excessive or inadequate heating could be minimized by independent control of RF currents and by varying the lengths of electrodes.

Microwave Heating. For superficial tumors, single-contact applicators operating at 433 MHz to 2450 MHz have been used. The shorter wavelength at these frequencies allows microwave radiation from a small applicator some focusing ability in tissues for selective hyperthermia. Because of the limited depth of energy penetration, these antennas have been applied to heating well-localized tumors extending to depths of up to 3 cm to 6 cm depending on the particular applicator (39,44,69,70).

The types of external applicators that have been reported for cancer hyperthermia include horns, mi­crostrip applicators, and circular, rectangular, and ridged waveguides (71,72,73,74,75,76,77,78,79). These ap­plicators are used with a high-permittivity, dielectric material to match them to tissue. In the case ofa water-like bolus, it serves also to provide surface cooling of the skin and to avoid the problem of burns and blisters. Mi­crostrip applicators are lightweight and have a low profile. They offer efficient energy coupling and are easier to use clinically (77,78,79). One limitation of a single applicator is its small area of tissue coverage. Another is that the SAR distribution cannot be modified during use, making it difficult to improve the nonuniform temperature distribution that are inevitably produced during patient treatments. One approach to overcome this problem is to scan the applicator over the tissues (80).

A favorable external system to treat tumors of wide area (tumors that exceed several cm in diameter) is the phased array consisted of multiple microstrip applicators. The primary advantage of the multielement array system is the ability to control electronically the SAR distribution by varying the amplitude and relative phase of each element, independently. Moreover, a planar or quasiplanar phased array operating at microwave frequencies can be used to improve depth of penetration for selective heating of deep-seated tumors in a variety of anatomic sites (81,82,83,84,85,86). A further advantage is that the SAR distribution can be adjusted during treatment, enabling it to enhance the homogeneity of temperature distribution in the target region. The added sophistication needed for controlling a multitude of array parameters is well within the capability of current electronic technology. Although a bolus of cooling fluids can be used to prevent undesirable heating of superficial tissues, tissue layers and curvatures in the near field of the applicator present considerable challenge to quality control in patient treatments. In practice, the commonly accepted SAR variation is 50% throughout the entire treatment region.

Intracavitary techniques can be used for certain tumors at hollow viscera and cavity sites such as the esophagus, cervix, bladder, prostate, and rectum (69,87). Properly designed intracavitary applicators and antennas can lead to a highly targeted heating of tumors and a reduced risk of unwanted heating of normal tissues. There are several reports of devices designed for various tumor sites (88,89,90,91). Clinical applications may require the antennas to be equipped with an integrated cooling system.

The technical difficulty in heating deep-seated tumors without overheating adjacent normal tissue con­fronted by external applicators has enabled interstitial array techniques to become a viable treatment modality (67,69,92,93,94,95). The technique has the capacity to adapt its SAR distribution to an irregularly shaped tumor volume and to provide uniform temperature in deep-seated tumors. Also in combination with brachytherapy, interstitial hyperthermia renders a treatment modality for malignancies with little additional risk to the patient (69,95,96).

The efficacy of interstitial microwave heat treatment of soft-tissue tumors is predicated on a sufficient temperature distribution throughout the tumor. A major determinant is the catheter antenna. Recent designs have provided microwave interstitial array systems capable of inducing uniform temperature distribution throughout the entire tumor volume without the need for insertion of the tip of the antenna well beyond the tumor boundary (97,98,99,100). That requirement was a major drawback of many older catheter antennas which had the tendency to produce a cold spot or low-heating zone near the distal tip of the antenna (67,101, 102,103), which creates an unnecessary situation for damage to normal tissue. A desirable feature of some of the newer catheter antenna designs, especially those with integral sleeves or coaxial chokes, is that the SAR distribution is independent of insertion depth (90,104,105). These antennas have also managed to alleviate the common problem of excessive heating of the skin from current accumulation at the insertion point. In the clinic, interstitial microwave antennas are inserted into plastic catheters implanted into the tumor. Computational and experimental studies have shown that SAR distributions vary with antenna design, catheter size and material, and air space between the antenna and the catheter (100,105,106).

Array configurations (i. e., geometry and antenna spacing) would also dictate the performance of the interstitial array treatment modality. Current microwave interstitial array systems rely mostly on equilateral triangle and square arrays of catheter antennas operating at 433 MHz to 2450 MHz and use element spacings of 10 mm to 20 mm. Theoretical and experimental results have shown that uniform power deposition and temperature distribution can be attained from both triangular and square arrays. However, power deposition and temperature elevation are higher for the triangular configuration at a given level of delivered microwave power. Moreover, for coherent phase excitations, constructive interference can provide SARs at the array centers an order of magnitude higher than those corresponding to a single interstitial microwave antenna. A flexibility afforded by an array of interstitial antennas is that the point of maximum SAR may be shifted from location to location by changing the amplitude and phase of each antenna. This would avoid low SAR spots during treatment and would ensure uniform tumor temperature over the entire treatment session. Nevertheless, it should be noted that ideal operating conditions are difficult to assure in the clinical setting.

MILLER EFFECT

While designing amplifiers, engineers may assume that the internal capacitances in the transistor are very small com­pared to the external capacitances. But in reality, capaci­tances do exist between the base and emitter (CBE) as well as between base and collector (CBC). This is shown in Fig. 4. It can be mathematically shown that the total input capacitance

CI = CBE + (1 + AV )(CBC)

In other words, the total input capacitance is the parallel combination of CBE and (1 + AV)CBC. The base-collector capac­itance has been amplified by a factor of 1 + AV. This is called the Miller effect.

As mentioned earlier, as the frequency increases, the value of the total input impedance decreases and thereby the fre-

Vc

MILLER EFFECT

Figure 4. Miller effect with the transistor internal capacitances CBC and Cbe.

MILLER EFFECT

Coupling Яс

MILLER EFFECT

MILLER EFFECT

VINO—— }|-

Cc

Emitter

by-pass

capacitor

Figure 5. The impedances of CC and CE are large at low frequencies, and portions of signal voltages may be lost.

quency response characteristics are affected. The Miller effect is especially pronounced with common-emitter amplifiers, be­cause they introduce a 180° phase shift between the input and the output. For example, the values of CBE and CBC may be small, say 5 pF and 4 pF. But when the transistor is used in an amplifier with a gain of 99, the total input capacitance will be large enough to affect the frequency output characteristics of the amplifier. This is because

ci — cbe + (1 + av )(cbc)

= 5 + (1 + 99)(4) — 405 pF

It is recalled that at low frequencies the coupling capacitor and the emitter bypass capacitors offer high impedances and therefore portions of signal voltage may be lost, as shown in

Fig. 5.

MILLER EFFECT

A’ 7

rj V feedback

^O, Miller — A1 — 1

The Miller effect is thus an extremely important concept in discussing feedback. Equations for calculating the Miller input impedance and Miller output impedance can be devel­oped, and are given below: