FOR MORE THAN 60 YEARS, therapeutic ultrasonography

Transcription

1 1318 ORIGINAL ARTICLE Variations in the Output Power and Surface Heating Effects of Transducers in Therapeutic Ultrasound Christian Kollmann, PhD, Gerda Vacariu, MD, Othmar Schuhfried, MD, Veronika Fialka-Moser, MD, Helmar Bergmann, PhD ABSTRACT. Kollmann C, Vacariu G, Schuhfried O, Fialka- Moser V, Bergmann H. Variations in the output power and surface heating effects of transducers in therapeutic ultrasound. Arch Phys Med Rehabil 2005;86: Objective: To determine the real emitted output power and maximum surface heating of commercial therapeutic ultrasound transducers emitting in air for various therapeutic regimens. Design: Surface temperatures of ultrasound transducers with frequencies of.05 to 3MHz were detected over 5 minutes by using a calibrated infrared thermographic camera; additionally, the indicated output power was checked with a radiation force balance. Setting: University center for biomedical engineering and physics and medical school for physical medicine and rehabilitation. Participants: Not applicable. Interventions: Not applicable. Main Outcome Measures: Power variations and surface temperatures of clinical devices were analyzed to determine whether they comply with obligatory limits given in International Electrotechnical Commission standard Results: Depending on the operation mode and the output power, surface temperatures ranged between 24.2 to 80 C within 5 minutes. Differences between measured and displayed power output (limit, 20%) ranged between 32% and 28%. Conclusions: The effectiveness of treatment is lowered if the value of emitted power is not known reliably. In the worst case, damage or irritation of the skin is possible, particularly in patients with sensory compromised skin. Damage may be caused by hot surfaces if the threshold level required to activate the device is lowered or if the device is defective. Improved thermal control units are necessary to prevent potential thermal hazards. Regular checks of transducer emission should be obligatory to ensure correct and precise function of the clinical devices. Key Words: Rehabilitation; Thermography; Transducer; Ultrasonography by American Congress of Rehabilitation Medicine and the American Academy of Physical Medicine and Rehabilitation From the Center for Biomedical Engineering and Physics (Kollmann, Bergmann) and Orthopaedical Hospital Speising (Vacariu), Vienna, Austria and Department of Physical Medicine and Rehabilitation (Schuhfried, Fialka-Moser), Medical University of Vienna; and Ludwig-Boltzmann-Institute of Nuclear Medicine (Bergmann), Vienna, Austria. No commercial party having a direct financial interest in the results of the research supporting this article has or will confer a benefit upon the author(s) or upon any organization with which the author(s) is/are associated. Reprint requests to Christian Kollmann, PhD, Center for Biomedical Engineering and Physics, Medical University of Vienna, Waehringer Guertel 18-20, A-1090 Vienna, Austria, christian.kollmann@meduniwien.ac.at /05/ $30.00/0 doi: /j.apmr FOR MORE THAN 60 YEARS, therapeutic ultrasonography (US) has been extensively applied in medical treatment to relieve acute and chronic pain, to accelerate tissue repair, and to treat musculoskeletal disorders. 1 Typical frequencies used for therapeutic US machines range from 0.8 to 3MHz in combination with variable intensities (I sata ) of.25 to 3.0W/cm 2. US devices with low-frequency ( 50kHz) and/or low-intensity output (.25W/cm 2 ) also are used. Use of these devices is predicated on the belief that the US wave s thermal effects (heating) and its nonthermal effects (caused by the pressure differences and particle displacement in a wave period) enhance therapeutic treatment and muscle repair. 2,3 The lowfrequency machines allow deep penetration because ultrasound beams are inherently less attenuated as the ultrasound wave frequency decreases, ignoring additional effects from wave collimation. The energy delivered by therapeutic US machines converts into heat in the human body. If the heat is not dissipated, localized tissue heating occurs, which results in a thermal therapeutic effect. According to Lehmann 4 and Castel, 5 an increase of tissue temperature by 1 C will result in a 13% increase of the metabolic rate. In general, moderate ultrasonic heating activates perfusion in the body region treated, 5-7 whereas a higher heat may decrease the viscoelastic properties of the collagenous tissues. 5-8 An ultrasound transducer for therapeutic applications usually consists of an air-backed piezoelectric element covered with a metallic layer. 9 A metallic layer having a thickness equal to the ultrasound wavelength ( ) or half the wavelength ( /2) provides both mechanical protection and coupling. This property sometimes combines with electromagnetic therapy properties. Through an alternating electric voltage applied to the piezoelectric element, the US wave is generated. A water-based gel or water layer is used to couple the transducer to the body and to ensure good contact. Three essential factors govern the output performance (ie, intensity, power, acoustic pressure) of the therapeutic US machine: (1) applied voltage amplitude, (2) the properties of the different tissue layers, and (3) the efficiency of the electrical-to-mechanical conversion process. This article focuses on the accuracy of the displayed power output settings of 4 ultrasound devices; it also compares maximum surface temperatures that occur if a transducer is not in contact with the body but is still emitting because of defective, changed, or missing contact detectors and longer timer settings. Possible consequences concerning therapeutic treatments are discussed. METHODS Measurements were performed for 4 clinically used, physiotherapeutic US devices from 4 different manufacturers. The devices operated in continuous-wave (CW) and/or pulsed-wave (PW) mode (table 1). The frequency range was between.05 and 3MHz. Three of the devices featured 2 separate transducers, one with a small effective radiating area (ERA) of 0.5 to 2.5cm 2 and another with a larger ERA of 4 to 12.8cm 2. In total,

2 EMISSION AND TRANSDUCER HEATING IN THERAPEUTIC ULTRASOUND, Kollmann 1319 Table 1: Types and Specifications of Therapeutic Ultrasound Machines Tested Model Impulsaphon MT 50 Sonoplus 590 UniPHY Phys-Assist Manufacturer and Year of Construction Dr. Born GmbH (Germany); 1996 Enraf-Nonius BV; (Netherlands); 1996 Phyaction Supporta (Netherlands); 1999 Orthosonics Ltd (UK); 2000 Type of Probe Frequency/ ERA Modes I SATA (W/cm 2 ) Circular flat 1MHz, 2.5cm 2 /5.0cm 2 Circular flat 1MHz/5.0cm 2 3MHz/0.5cm 2 Circular flat 0.7 1MHz, cm 2 CW PW CW 16/48/100Hz CW PW 1:2/1:4/1: Spherical CW 5/15/30/50mW/cm 2.046MHz, 12.8cm 2 NOTE. 16Hz is 16ms pulse, 47ms pause; 48Hz is 4ms pulse, 16ms pause; 100Hz is 2ms pulse, 8ms pause. 1:2 (PW mode) is 1 pulse period followed by 2 pause periods (others analog). Abbreviations: ERA, effective radiating area; I SATA, spatial averaged time/averaged intensity, nominal. we investigated 41 different therapeutically used operating modes (PW/CW) and intensities. All transducers were carefully cleaned of gel and the surface dried before performing the measurements. An infrared (IR)- thermography camera (Thermovision 900 TE) a (spectral response, m; image size, pixels; 12 bit) was positioned perpendicularly above the transducer, which was fixed in a rig. The optical focus of the camera was adjusted to be directly on the surface of the transducer at a distance of 80cm (spatial resolution, 1.7mrad; 20 SW lens). The camera s thermal resolution at 23 to 80 C was 0.1 C with an accuracy of 1 C. We compensated for the ambient temperature (24 1 C) and calculated the emissivity of the particular transducer surface by using the ambient temperature before each measurement. The camera was connected to a personal computer, which enabled its functions to be handled by software (ThermoCam Researcher 2000). a Thermal images were acquired when the transducer was operating in air for 5 minutes. Its surface temperature was imaged every 5 seconds and stored digitally as a pseudocolor image (fig 1). After recording a complete IR-image sequence, the camera software recorded the values within the effective area of the transducer. In an off-line process, the pseudocolor-coded temperature values were automatically recoded into numeric values for the total sequence and were stored. Normally, therapeutic transducers do not emit into air because of their automatic sensors, which prevent transmission until contact is detected. One must apply contact gel or another thin film layer to properly transmit the US energy into the body. The transducer can emit unintentionally, however, if its detection threshold is lowered, defective, not implemented, or is exceeded by contact gel adhering to transducer surface. To simulate and evaluate these worst-case situations, a thin (25 m) Saran foil b covering a film of gel was applied to the transducer surface to match the impedance load condition for operation mode. 9 To measure the acoustical output power (p) of each machine, we used a calibrated radiation force balance c with an uncertainty of 3% up to 30W at resolution in steps of 2mW. The difference (in percent) between measured and displayed output power was calculated by using the formula 10 : power measured power displayed difference(%) power displayed 100 (1) RESULTS Maximum surface temperatures ranged from 24.2 up to 64.0 C at 5 minutes for different setup configurations (figs 2, 3). Figure 2 shows the rise in surface temperature over time for equipment operating in CW mode. A transducer with a large ERA typically generates a higher surface temperature at 5 minutes than does a transducer with a smaller ERA on the same machine, operating with the same nominal intensity. One machine showed at 4 minutes a maximum surface temperature of around 82 C by using the highest intensity available at CW mode and a large-transducer ERA. Figure 3 shows the rise in Fig 1. Examples of IR-thermography images showing transducer surfaces operating in CW mode: (A) circular flat transducer driven with an intensity of 2W/cm 2 and (B) spherical transducer driven with an intensity of 50mW/cm 2.

3 1320 EMISSION AND TRANSDUCER HEATING IN THERAPEUTIC ULTRASOUND, Kollmann Fig 2. Increase of transducer surface temperature for equipment operating in CW mode with different intensities (a small ERA transducer for the same machine is indicated with [s] as seconds within the legend). temperature for the machines operating in PW mode. The surface temperatures reached after 5 minutes are lower than for CW mode. The surface temperature T surf can be fitted well (see drawn lines in figs 2, 3) by using the exponential function (T surf T RT a(1 exp[ b t]), where t is the time in seconds, T RT the room temperature in degrees Celsius, and a and b are coefficients depending on the specific transducer. For most operating modes, the transducer surface reached its highest temperature (saturation effect) within 5 minutes after being switched on. A surprising result was the high increase of the surface temperature for the low-intensity low-frequency machine (see fig 2, Phys-Assist d ). Data sets from 2 machines (Impulsaphon, e UniPHY f ) could not be measured over the full period: in both cases, the maximum output settings (CW mode, maximum intensity) caused such high absolute surface temperatures (51 C) within 1 minute, 82 o C within 4 minutes operating time, respectively, that the thermal control unit of the equipment switched off the device. In figures 2 and 3, the limit of maximum surface temperature allowed by an international standard for diagnostic imaging ultrasound transducers 11 is indicated as a broken horizontal line. The standard requires that the maximum surface temperature for these kinds of transducers should not exceed 50 C when applied to patients. Figures 4 and 5 compare the displayed output powers for the various machine settings with the actual measured power they emit. It is remarkable that the majority of the machines emit considerably less power than that displayed on screen. Only 1 transducer (UniPHY ERA, 1cm 2 ) emits more power than displayed at all operation modes. The differences ranged between 32% and 28%. The output power for 1 machine (Phys-Assist) could not be measured because the emitted frequency was outside the working range of the radiation force balance. DISCUSSION During operation of the transducer in air, each of the layers absorbs energy that is converted into heat and, in total, the transducer surface warms up. The reason is the large difference between the acoustical impedances of air (Z air, 0.4MRayl) and the metallic layer of the transducer (Z metal, 35MRayl). 9 The amount of transmitted (I SATA[T] ) and reflected (I SATA[R] ) intensity of the initial intensity (I SATA[E] ) can be calculated, respectively, according to the following 2 formulas: 4Z metal Z air I SATA(T) I SATA(E) (2) Z metal Z air 2 I SATA(R) I SATA(E) Z 2 metal Z Z metal Z air (3) Most ( 95%) of the ultrasonic energy (acoustical power) remains inside the transducer. The impedances Z air and Z metal shown previously are a typical example of this phenomenon and can produce the shown high temperatures, depending on the operating mode and setup configuration that are used. The maximum temperature that can be reached within normal application times is not known, nor do we know the dependence of temperature on the machine s settings. The settings used in

4 EMISSION AND TRANSDUCER HEATING IN THERAPEUTIC ULTRASOUND, Kollmann 1321 Fig 3. Increase of transducer surface temperature for equipment operating in PW mode with different intensities (a small ERA transducer for the same machine is indicated with [s] within the legend). the present study were representative of those used routinely in physiotherapeutic treatment regimens for acute or chronic pain. Our measurements show that very high temperatures can be reached within a few minutes, not only for transducers operating in CW mode but also for those operating in PW mode. Another important outcome of this study was the relatively Fig 4. Displayed versus measured power for various ERA of machines operating in CW mode. high surface temperature of transducers from the low-intensity low-frequency machines. Usually, it is assumed that an output intensity (I SATA p/ ERA) of around 50mW/cm 2 is too low to produce a noticeable increase in transducer surface temperature. According to our findings, this assumption is not true. Therefore, this type of machine should be used carefully, too. Thermal Transducer Regulation The study revealed that the thermal control function in some machines turns the power off too late (ie, temperatures rise too high above the limit given by the International Electrotechnical Commission [IEC] 11 ). Some devices do not have a thermal control and their automatic contact sensor may be missing or defective. Output display accuracy. It was surprising that the machines typically generate considerably less acoustical power than displayed (figs 4, 5). In general, choosing a higher output level will produce a more inaccurate display than choosing a lower output level. Some (24.3%) of the power differences exceed the 20% tolerance level given in IEC standard or by the Food and Drug Administration, 13 which indicates that maintenance and calibration are needed. These high-output variations were reported internationally for other machines also; another group 14 detected intolerable output differences in 37.8% of the checked equipment. Smaller variations in power, internal mechanisms for self-adjusting power levels, and regular calibra-

5 1322 EMISSION AND TRANSDUCER HEATING IN THERAPEUTIC ULTRASOUND, Kollmann Fig 5. Displayed versus measured power for various ERA of machines operating in PW mode. tion checks are needed to ensure more reliable therapeutic equipment for treatment. Clinical impact. Applying a hot transducer surface to the skin can negatively affect medical treatment: patients can be frightened or, in the worst case, the skin can be harmed by the local thermal impact. 15 Knowledge of transducer surface temperatures is important for improving simulation models too. These models predict the temperature increase within tissue and show the highest increase on skin surfaces 16 for the same reason: the large impedance difference between transducer and skin (Z skin, 1.76MRayl; I SATA(R), 82%). The models presented in Koch 16 were calculated for diagnostic ultrasound machines having lowintensity output, but the impact of the temperature increase at skin surface will be higher if therapeutic machines are used because of their.64 to 40W acoustical power. Figure 6 shows skin heating of the inner forearm for 2 different transducer regimens operating with a CW emission of 1W/cm 2 (figs 6B, 6C), along with a reference image (fig 6A). A preexisting hot surface of a transducer causes a larger area of high heat on the skin s surface (fig 6B) than that found during a normal therapeutic regimen (fig 6C). But even the normal therapeutic application heats up small areas of the skin surface locally, with hot spot temperatures that could destroy tissue if the treatment duration is long and the intensity is high or if the regimen is static. This localized excessive heating could become a problem especially for patients with sensory disturbances. In addition, the transducer may be damaged by frequent cycles of high temperature differences. The junction between the piezoceramic element and the covering layers may fail, which will shorten the transducer s life. 9 The efficiency of treatment is decreased, too, by lowering the depth where the highest peak temperatures in tissue are generated. 4 The discrepancy between displayed and actual power or intensity emitted by the machines (see figs 4, 5) is another drawback in high-power applications. Because a specific increase in temperature is needed to achieve a temperature-mediated therapeutic effect, low power would be a problem in treating deeply situated structures. The temperature increase for deeper tissue layers would be lower if the output power were effectively 20% to 30% lower than that shown at the display. 17 If appropriate heating is dependent on frequency an assumption that has not been verified 18 longer treatment durations and changes of the treatment protocol would be needed to obtain the therapeutic effect sought with the higher power levels. CONCLUSIONS Therapeutic US machines emitting without coupling are able to heat up the transducer surface within a few minutes to temperatures that, if applied to a patient, could cause skin irritations or, in the worst case, local skin burns. A hot transducer surface can develop if the sensors that are supposed to detect missing or changed contact are set to a lower threshold. It can also develop if the machine s thermal control systems or timer should fail or if damage from unintentional continuous operation should occur. To avoid these incidents, we recommend that users actively switch off the power, ensure that the transducer is completely clean of gel after treatment, and verify that the device s contact detection threshold does not change from that implemented by the manufacturer. For frequently used equipment, one should provide enough time between patients to allow the transducer surface to cool to ambient temperature. This action will help to achieve a high level of uniform therapeutic heating at the intended body depth. 4 If these actions cannot be implemented, it would be advisable at least to use an additional layer (eg, water bag) between the patient s skin and the transducer, or to conduct the treatment under water, if possible.

6 EMISSION AND TRANSDUCER HEATING IN THERAPEUTIC ULTRASOUND, Kollmann 1323 Fig 6. IR-thermography images of the inner forearm after 1 minute of transmission in CW mode at 1W/cm 2 by the same transducer under 3 conditions: (A) the transducer was statically coupled to the skin, but not activated (reference temperature); (B) the transducer was preheated, emitting in air for 5 minutes before being applied to the skin (surface temperature, 55 C); and (C) the transducer was applied at ambient surface temperature then operated. Stricter international variation limits are needed regarding the power emission. 11,12 Also needed are an obligatory thermal control and self-adjusting constant-power emission for each machine. Finally, recurrent output controls are advised to inform the user about the status of the machine and about the necessary recalibration. Acknowledgments: We thank all contributors to this project, including Mrs. Laikie and her team of physiotherapists, Monika Knötig and Monika Putzer, for preparing and allocating the devices and for providing us with additional therapeutic information. References 1. Robertson VJ, Baker KG. A review of physiotherapeutic ultrasound: effectiveness studies. Phys Ther 2001;81: Rubin C, Bolander M, Ryaby JP, et al. The use of low-intensity ultrasound to accelerate the healing of fractures. J Bone Joint Surg Am 2001;83: Dyson M, Preston R, Woledge R, et al. Long-wave ultrasound. Physiotherapy 1999;85: Lehmann JF. Ultrasound therapy in therapeutic heat and cold. 3rd ed. London: Licht; Castel JC. Therapeutic ultrasound. Rehabil Ther Products Rev 1993;1-2: Draper DO, Castel JC, Castel D. Rate of temperature increase in human muscle during 1 MHz and 3 MHz continuous ultrasound. J Orthop Sports Phys Ther 1995;22: Maxwell L. Therapeutic ultrasound: its effects on the cellular and molecular mechanisms of inflammation and repair. Physiotherapy 1992;78: Kottke FJ, Lehmann JF, editors. Krusen s handbook of physical medicine and rehabilitation. 4th ed. Philadelphia: WB Saunders; Moreno E, Gonzalez G, Leija L, et al. Performance analysis of ultrasono-therapy treatment head with contact detection. IEEE Trans Ultrason Ferroelectr Freq Control 2003;50: Pye S, Zeqiri B. Guidelines for the testing and calibration of physiotherapy ultrasound machines. York: Institute of Physics and Engineering in Medicine; Report No International Electrotechnical Commission. Medical electrical equipment. Part 2-37: Particular requirements for the safety of ultrasonic medical diagnostic equipment. Chicago: IEC; Publication No International Electrotechnical Commission. Medical electrical equipment. Part 2: Particular requirements for the safety of ultrasonic therapy equipment. 2nd ed. Chicago: IEC; Publication No US Food and Drug Administration. 21 CFR (1994). 14. Daniel DM, Ruprecht RL. Calibration and electrical safety status of therapeutic ultrasound used by chiropractic physicians. J Manipulative Physiol Ther 2003;26: National Council on Radiation Protection and Measurements. Exposure criteria for medical diagnostic ultrasound: I. Criteria

7 1324 EMISSION AND TRANSDUCER HEATING IN THERAPEUTIC ULTRASOUND, Kollmann based on thermal mechanisms. Bethesda: NCRP; p Report No Koch C. [Thermal effects of ultrasound] [German]. Ultraschall Med 2001;22: Cambier D, D Herde K, Witvrouw E, et al. Therapeutic ultrasound: temperature increase at different depths by different modes in a human cadaver. J Rehabil Med 2001;33: Demmink JH, Helders PJ, Hobaek H, et al. The variation of heating depth with therapeutic ultra-sound frequency in physiotherapy. Ultrasound Med Biol 2003;29: Suppliers a. FLIR Systems Inc, 25 Esquire Rd, North Billerica, MA b. Dow Chemical, 2030 Dow Ctr, Midland, MI c. Model UPM-DT-1; Ohmic Instruments Co, 508 August St, Easton, MD d. Orthosonics Ltd, Bremridge House, Bremridge, Ashburton, Devon TQ13 7JX, UK. e. Impulsaphon, Dr. Born GmbH Ultraschall, Frankfurt, Germany. f. Phyaction Supporta, Uniphy Apparaturen B.V., Eindhoven, The Netherlands.