Many medical devices serve the portable healthcare and emergency response markets. Examples of these devices are heart-rate monitors, glucometers, electrocardiogram (ECG) monitors, ultrasound imaging devices, and diagnostic medical imaging devices such as digital radiographic detectors. Regardless of the application, these devices must be small in size, lightweight and battery powered to provide the user of the device with optimum mobility and ease of use. However, the requirements necessary to achieve portability result in severe constraints on space, weight and power dissipation causing an increase in the amount of heat energy generated by the components of the medical device. The primary sources of heat are the various integrated circuit components, and rechargeable batteries that power the device when in use or when battery charging takes place. The net result is that the heat contributes to an overall rise in temperature with both application and structural effects on the medical device. Structural effects or excessive heat generated by small portable electronic devices reduces battery life, reduces component life, reduces the reliability of the device, and increases device failure.
In diagnostic medical imaging devices the problem of heat generation is a greater concern due to high power requirements, usage of complex circuitry for optimal performance that is highly sensitive to heat, and patient safety. In particular, while the high power and high circuit density required by the portable battery powered diagnostic medical imaging devices further exacerbate the problem of heat generation, these devices must satisfy certain medical safety requirements regulating the maximum external surface temperature of the device to insure patient safety. Present medical safety requirements regulating temperature mandate that the maximum allowable external surface temperature of a medical device (i.e., the “skin” temperature) not exceed 50 degrees ° C. (122 degrees Fahrenheit), thereby ensuring that contact with a patient will not result in patient discomfort or burning. More specifically, there several regulations and rules regarding the temperature of medical devices, such as IEC 60601-1 promulgated by the International Electrotechnical Commission. These regulations are known to the practitioners of the art.
In the case of digital radiographic or digital x-ray, the electronics in the detector generates a significant amount of heat during image acquisition, due to their electrical power consumption, but can be operated at reduced power when no image is being taken. These devices include a source for projecting an x-ray beam toward an object to be analyzed, such as a medical patient. After the beam passes through the patient, an image intensifier converts the radiation into a signal. With solid state digital x-ray detectors, the photodiode detector elements produce electrical signals that correspond to the brightness of the picture element in the x-ray image projected onto the detector. The signals from the detector elements are read out individually and digitized for further image processing, storage and display, typically by a computer. However, to achieve the required image quality, some time is required for electronic signal levels to fully stabilize between the image detector being restored to full power and acquisition of an image. This stabilization time to interfere with the process of acquiring the image is undesirable because the patient may be in an uncomfortable position, required to hold their breath for the image, or other reasons.
To add to the complexity of the problem the imaging performance characteristics of the detector vary with the temperature of the panel and the temperature of the pixel array. For optimum imaging performance, the panel temperature must remain within a range of temperatures. Techniques utilizing higher x-ray power and longer exposures are in demand in order to provide better images. Thus, there is an increasing demand to remove as much heat as possible from the x-ray tube, as quickly as possible, in order to increase the x-ray exposure power and duration before reaching the operational limits of the tube. At full power, the electronics of the detector consume sufficient power and generates sufficient heat to require a thermal management control subsystem to maintain the panel within the optimum imaging temperature range. Previous attempts at developing cooling systems to remove the heat energy from the relatively high-density packaging of radiographic digital image detectors have primarily used thermal convection systems. These systems move large volumes of heat absorbing air or fluid through the radiographic digital image detector to remove the heat energy created by operation of the device. This large volume requires open spaces around the digital radiographic detector. The necessity of these open spaces limits the overall density of the storage devices relative to the volumetric space of the storage system. However, this technique depletes the finite amount of energy especially in a portable device.
For the reasons stated above, and for other reasons stated below which will become apparent to those skilled in the art upon reading and understanding the present specification, there is a need in the art for a reliable, simple and efficient manner to provide a thermal management system in a portable battery powered electronic device, and particularly, in a portable battery powered diagnostic medical imaging device, which addresses the foregoing problems. There is also a need for improved management of power consumption in portable devices such as digital radiographic detector to increase power conservation and increase efficiency.
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