Abstract:
Infusion pumps according to the present invention maximize fluid throughput while minimizing vaporization of gas by employing specific flow path architecture, flow path dimensional ranges, and voltage and frequency ranges for activation of piezoelectric bodies.

Description:
RELATED APPLICATIONS 
     This application claims priority to U.S. Provisional Application Ser. No. 61/287,881 filed Dec. 18, 2009, entitled MEMS Pump for Medical Infusion Pump; U.S. Provisional Application Ser. No. 61/287,903 filed Dec. 18, 2009, entitled Pump Stay; U.S. Provisional Application Ser. No. 61/287,912 filed Dec. 18, 2009, entitled Micro Infusion Pump System Software; U.S. Provisional Application Ser. No. 61/287,991 filed Dec. 18, 2009, entitled Central Venous Pressure Monitoring Using Micro Infusion Pump, the contents of which are each incorporated in their entirety herein. 
    
    
     FIELD OF THE INVENTION 
     The present invention relates to medical infusion pumps and related methods and, more particularly, to infusion pumps employing the piezoelectric effect for medical and healthcare related applications. 
     BACKGROUND OF THE INVENTION 
     Fluid pumps can be driven based on various design principles including the piezoelectric effect. The piezoelectric effect can be employed to indirectly cause fluid flow, for example a piezoelectric driven motor or actuator can be used to linearly displace a plunger to push fluid from a reservoir or to rotate a rotor in a peristaltic-type pump. For example, U.S. Publication Nos. 2009/0124994 to Roe and 2009/0105650 to Wiegel et al., and U.S. Pat. Nos. 7,592,740 to Roe, and 6,102,678 to Perclat teach the application of such technologies to infusion pumps used in the medical and health care industries. 
     Alternatively, the piezoelectric effect can be employed to cause fluid flow through the direct manipulation of a fluid chamber or flow path, for example through vibration of an internal surface of a fluid chamber. Such microelectromechanical system, or MEMS, micropumps can be fabricated using known integrated circuit fabrication methods and technologies. For example, using integrated circuit manufacturing fabrication techniques, small channels can be formed on the surface of silicon wafers. By attaching a thin piece of material, such as glass, on the surface of the processed silicon wafer, flow paths and fluid chambers can be formed from the channels and chambers. A layer of piezoelectric material, or a piezoelectric body such as quartz, is then attached to the glass on the side opposite the silicon wafer. When a voltage is applied to the piezoelectric body, a reverse piezoelectric effect, or vibration, is generated by the piezoelectric body and transmitted through the glass to the fluid in the chambers. In turn, a resonance is produced in the fluid in the chambers of the silicon wafer. Through the inclusions of valves and other design features in the fluid flow paths, a net directional flow of fluid through the chambers formed by the silicon wafer and the glass covering can be achieved. 
     MEMS micropumps have become an established technology in the inkjet printer industry. Technological developments relating to increased definition and ink throughput for piezoelectric micropumps, or MEMS micropumps, for inkjet printers have achieved more precise printing with smaller ink throughputs. For example, it has become possible to control the ink throughput of inkjet printers employing MEMS micropumps at the picoliter level. Furthermore, in order to address the problems associated with uneven printing in inkjet printers due to the vaporization of gas dissolved in the ink, considerable development has also been directed to providing inkjet printers with structures for degassing the ink. 
     MEMS micropumps employing the piezoelectric effect have also been contemplated for use in small and large-volume infusion pumps, i.e. pump systems that are typically employed to infuse fluids, medications, and nutrients into a patient&#39;s circulatory system. For example, with respect to small-volume infusion systems, U.S. Pat. Nos. 3,963,380 to Thomas, Jr. et al.; 4,596,575 to Rosenberg; 4,938,742 to Smits; 4,944,659 to Labbe et al.; 5,984,894 to Poulsen et al.; and 7,601,148 to Keller all describe various micropumps intended for implantation into a patient in order to administer small amounts of pharmaceuticals, such as insulin. Similarly, U.S. Publication No. 2007/0270748 to Dacquay et al. describes a piezoelectric micropump integrated into the tip of a syringe for very low volume delivery of ophthalmic pharmaceuticals to a patient&#39;s eye. 
     In contrast to inkjet printers and small-volume infusion micropumps, large-volume infusion pumps must be operable to provide significantly increased fluid throughput. However, as fluid throughput, or fluid flow rates are increased, the potential for the vaporization of dissolved gas correspondingly increases. Those skilled in the art will recognize that the vaporization of dissolved gas within the fluid flow paths of infusion pump systems presents a significant health hazard to patients receiving infusions. While the problems associated with the vaporizations of dissolved gas in inkjet printer micropumps, systems in which fluid throughputs are relatively low, has largely been addressed through the development of degassing technologies, satisfactory solutions have not been presented for high-throughput micropumps, such as infusion pumps, used in the health and medical industry. U.S. Publication No. 2006/0264829 to Donaldson and U.S. Pat. No. 5,205,819 to Ross et al. described large-volume infusion systems employing piezoelectric micropumps; however, neither of these systems provides solutions directed to overcoming the problems associated with vaporization of dissolved gas at high fluid throughputs. 
     What is needed in the field is a highly accurate infusion pump system that provides high fluid throughput while reducing or eliminating the risk of the vaporization of dissolved gasses within the fluid flow path. 
     OBJECTS AND SUMMARY OF THE INVENTION 
     The infusion pumps according to the present invention provide highly accurate infusion pump systems that provides high fluid throughput while reducing or eliminating the risk of vaporization of dissolved gasses within the fluid flow path. Infusion pumps according to the present invention achieve these advances by employing specific flow path architecture, flow path dimensional ranges, and voltage and frequencies ranges that, relative to one another, serve to provide an infusion pump that maximize fluid throughput while minimizing vaporization of gas. 
    
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
       These and other aspects, features and advantages of which embodiments of the invention are capable of will be apparent and elucidated from the following description of embodiments of the present invention, reference being made to the accompanying drawings, in which: 
         FIG. 1  is a perspective view of an infusion pump according to one embodiment of the present invention. 
         FIG. 2   a  is a perspective view of a micropump according to one embodiment of the present invention. 
         FIG. 2   b  is a cross-sectional view taken along line b of  FIG. 2   a  of a micropump according to one embodiment of the present invention. 
         FIG. 2   c  is a plan view of a pump base of a micropump according to one embodiment of the present invention. 
         FIG. 2   d  is an expanded plan view of a portion  35  of  FIG. 2   c  of a pump base of a micropump according to one embodiment of the present invention. 
         FIG. 3  is a plan view of flow channels and cambers of a micropump according to one embodiment of the present invention. 
         FIG. 4  is a graphical representation of a power provided to a micropump according to one embodiment of the present invention. 
         FIGS. 5   a  and  5   b  are cross-sectional views of a micropump according to one embodiment of the present invention. 
         FIG. 6  is a plan view of a micropump according to one embodiment of the present invention. 
         FIG. 7  is a plan view of a micropump according to one embodiment of the present invention. 
         FIG. 8   a  is a side elevation view of a micropump according to one embodiment of the present invention. 
         FIG. 8   b  is a plan view of a micropump according to one embodiment of the present invention. 
     
    
    
     DESCRIPTION OF EMBODIMENTS 
     Specific embodiments of the invention will now be described with reference to the accompanying drawings. This invention may, however, be embodied in many different forms and should not be construed as limited to the embodiments set forth herein; rather, these embodiments are provided so that this disclosure will be thorough and complete, and will fully convey the scope of the invention to those skilled in the art. The terminology used in the detailed description of the embodiments illustrated in the accompanying drawings is not intended to be limiting of the invention. In the drawings, like numbers refer to like elements. 
     Infusion pumps according to the present invention provides unique flow path and fluid chamber structure in order to provide highly accurate fluid flow and relatively high fluid throughput while reducing or eliminating the risk of vaporization of dissolved gas within the fluid flow path. 
     As shown in  FIG. 1 , an infusion pump  10  according to one embodiment of the present invention comprises a pump body  20   a  having an inlet port  22 , an outlet port  24 , and a communication port  26 . Situated within body  20   a  is a MEMS micropump. As shown in  FIG. 2   a  and the cross-sectional view of  FIG. 2   b  taken along line b of  FIG. 2   a , the MEMS micropump, or micropump,  30  comprises a piezoelectric body  32  having a thickness indicated as d 3  attached to one side of a vibration layer  36 . The vibration layer  36  may be fabricated from glass, quarts, alloy or a polymer-based material. An opposite side of the piezoelectric body  32  is in electrical contact with an electrode  34 , and an opposite side of the vibration layer  36  is affixed to a pump base  38 . As shown in  FIG. 2   b  and the plan views of the pump base  38  shown in  FIGS. 2   c  and  2   d , formed within the pump base  38  are flow channels and fluid chambers. More particularly, the pump base  38  comprises inlet channel  40 , outlet channel  42 , and chamber  44 . The chamber  44  has a length I 1  and a width I 2 . A distance from a bottom  46  of the chamber  44  to a top of the chamber  44  formed by the vibration layer  36  is indicated as a depth d. It will be understood that the depth d is determined in the approximate center or middle of the chamber  44 . An inclination or angle of the bottom  46  of the chamber  44  relative to the bottom  48  of the pump base  38  is indicated as an angle R 1 . While only one angle R 1  has been described, it will be understood that the bottom  46  of the chamber  44  may be angled relative to the bottom  48  of the pump base  38  in more than one plane of reference. As shown in  FIG. 2   b , as a result of the inclination or angle of the bottom  46  of the chamber  44 , the chamber  44  has a cross-sectional area that is greatest approximate the outlet channel  42 . The inlet channel  40  and the outlet channel  42  are each formed on opposite sidewalls of the chamber  44 , are each continuously open to fluid flow, and each comprises a differently configured channel constriction or cross-sectional shape. 
     The piezoelectric body  32  may comprise various known materials having piezoelectric properties, including naturally occurring crystals such as quartz, man-made crystals and man-made ceramics and polymers. 
     The pump base  38  may be fabricated from silicon; such as the silicon wafers employed in known integrated circuit fabrication techniques, or may be fabricated from other sufficiently rigid materials such as various metals, alloys, and polymers. It will be understood that, based upon the specific material(s) from which the pump base  38  is fabricated as well as, the specific technique employed for creating the channels and chambers within the pump base  38 , the surfaces of the channels and chambers may be smooth and/or stepped. For example, if the pump base  38  is fabricated using known integrated circuit fabrication techniques, the inclination or angle of the bottom  46  of the chamber  44  may be formed by removing silicon in defined steps according to the various masks employed. Accordingly, the sloped surface of the bottom  46  of the chamber  44  would be stepped rather than smooth. Conversely, if the pump base  38  is fabricated from a metal or alloy, the inclination or angle of the bottom  46  of the chamber  44  may be formed such that the surface of the bottom  46  of the chamber  44  is smooth. 
     Formed within the inlet channel  40  is a constriction A, and formed within the outlet channel  42  is a constriction B. It will be noted that the constrictions A and B may narrow in a horizontal, vertical, or horizontal and vertical directions. A length of the inlet channel  40  from the constriction A to the side  50  of the chamber is indicated as I 3 . A width of the inlet channel  40  upstream of the constriction A is indicated as L 1 , and a width of the constriction A is indicated as L 2 . 
     As shown best in  FIG. 2   d , between the constriction A and a chamber inlet  52 , the inlet channel  40  has a funnel-like or triangular form. An angle defined by this funnel-like form of inlet channel  40  between the constriction A and a chamber inlet  52  is indicated as R 2 . The width of the chamber inlet  52  is indicated as L 3 . In operation, the piezoelectric body  32  vibrates in the directions indicated by arrow  33  as a result of a voltage V applied to the piezoelectric body  32  through the electrode  34  at a frequency n. The vibration of the piezoelectric body  32  is transferred to a fluid within the chamber  44  through the vibration layer  36  thereby producing a resonance in the fluid within the chamber  44 . A pumping, or a net directional flow, of the fluid within the chamber  44  occurs in accordance with the resistances created by the constrictions A and B formed in the inlet channel  40  and the outlet channel  42  respectively. 
     It will be understood that the resonance created within the chamber  44 , and thereby the flow of fluid through the micropump  30 , is influenced by geometry of the chamber  44  and inlet and outlet channels  40  and  42 . Also influencing the resonance is the frequency, the voltage, and the shape of the wave that contribute to the piezoelectric effect of the piezoelectric body. Therefore, micropumps  30  according to the present invention are optimized with respect to the following parameters: (1) the length I 1  of the chamber  44 ; (2) the width I 2  of the chamber  44 ; (3) the depth d of the chamber  44 ; (4) the angle R 1  defined by the bottom  46  of the chamber  44  relative to the bottom  48  of the pump base  38 ; (5) the number of micropumps or micropump chambers  44  employed in the system; (6) the configuration of the constriction A relevant to the inlet channel  40  up stream of the constriction A; (7) the configuration of the constriction A relevant to the chamber inlet  52  downstream of the constriction A; vibration layer  36  (8) the thickness d 3  of the piezoelectric body  32  and the number of piezoelectric bodies  32  employed; and (9) the voltage V applied to the piezoelectric body  32  at the frequency n. 
     Turning first to the relevant parameters regarding the chamber  44  of the micropump  30 , in order to achieve a high fluid throughput, micropumps  30  according to the present invention employ a relatively large-volume chamber  44 . The shape of the chamber  44 , that is to say, the relationship of the length to width to depth of the chamber  44  is an important consideration in order to optimize the resonance phenomena generated in the fluid within the chamber  44 . For example, the depth d of the chamber  44  influences the magnitude of the resonance which, in turn, increases fluid throughput. However, at extreme magnitudes, the resonance of the liquid in the chamber  44  will undesirably occur only at the upper layer of the fluid and fluid throughput will decrease. 
     Relative to the other optimized parameters herein provided for micropumps  30  according to the present invention, the distance d from a bottom  46  of the chamber  44  to a top of the chamber  44  formed by the vibration layer  36  is preferably 50 to 300 micrometers and more preferably 100 to 200 micrometers. The length I 1  of the chamber  44  is preferably within the range of 1 to 30 millimeters, and the width I 2  of the chamber  44  is preferably within the range of 1 to 5 millimeters. Within the preferred range on depths d, a preferred range of length to width ratios is 1:1 to 6:1 and more preferably 4:1. For example, an optimized chamber  44  may have a length I 1  of 4 millimeters and a width I 2  of 1 millimeters and a depth of 200 micrometers. 
     Also related to the shape of the chamber  44  is the angle R 1  defined by the bottom  46  of the chamber  44  relative to the bottom  48  of the pump base  38 . The angle R 1  serves to bias the resonance towards the outlet channel  42  side of the chamber  44 , i.e. towards the right side of the chamber  44  as shown in  FIGS. 2   b  and  2   c . While a greater angle R 1  increases the fluid throughput, at overly steep angles, the resonance at the downstream side of the chamber  44  and thereby fluid throughput will tend to decrease. Relative to the other optimized parameters herein provided for micropumps  30  according to the present invention, the angle R 1  defined by the bottom  46  of the chamber  44  relative to the bottom  48  of the pump base  38  is preferably in the range of zero to 50 degrees and more preferably between 0.01 to 10 degrees. 
     In situations in which one micropump  30  does not provide the desired throughput, a micropump  30  having increased throughput can be formed by combining a number N of chambers  44  in parallel or in series. In such multi-chamber  44  pumps, identical chambers  44  may be combined, or as shown in  FIG. 3 , chambers  44  having differing sizes and throughputs may be combined. Pumps having a plurality of chambers  44  of different sizes and or dimensions operable of providing different fluid throughputs, allow for a more precise throughput control and a broader throughput range. For example, in cases such as those wherein the throughput is to be varied, advantages such as better precision and a wider range of throughputs can be achieved by combining a 100 milliliter per hour chamber  44   a  and a 10 milliliter per hour chamber  44   b . In medical settings, this is advantageous in so much as it is not necessary to change pumps, even when changing from 100 milliliter per hour of a drug to 10 milliliter per hour of a drug. 
     Turning now to the relevant parameters regarding the constriction A and the inlet channel  40  upstream of the constriction A. It is noted that a smaller width L 2  of the constriction A is associated with increased throughput, however, a smaller width L 2  is also associated with increased negative pressure downstream from the constriction A. This, in turn, results in an increased undesirable vaporization of dissolved gas. Relative to the other optimized parameters herein provided for pumps according to the present invention, the width L 2  of the constriction A is preferably in 30 to 200 micrometers and more preferably 40 to 80 micrometers, for example 50 micrometers. The width L 1  of the inlet channel  40  upstream of the constriction A is preferably 50 to 300 micrometers and more preferably 80 to 100 micrometers, for example 80 micrometers. The range of the ratio L 2 :L 1  is preferably 0.13 to 0.67. 
     Turning now to the relevant parameters regarding the constriction A and the inlet channel  40  downstream of the constriction A. The constriction width angle R 2  is defined by this funnel-like, or triangular, form of inlet channel  40  between the constriction A and the chamber inlet  52 . Alternatively stated, the constriction width angle R 2  refers to the angle formed at the apex of a triangle formed by the length I 3  and the width L 3 . While increasing the constriction width angle R 2  achieves a greater fluid throughput, increasing the constriction width angle R 2  also increases the negative pressure upstream of the constriction A. The angle R 2  is determined by the formula: 
     
       
         
           
             
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     Alternatively stated the angle R 2  is equal to two times the tangent of L 3  minus L 2  times 0.5 divided by I 3 . Relative to the other optimized parameters herein provided for pumps according to the present invention, width L 2  of the constriction A is preferably in 30 to 200 micrometers and more preferably 50 micrometers; width L 3  of the chamber inlet  52  is preferably 50 to 300 micrometers and more preferably 80 micrometers; and the length I 3  of the inlet channel  40  from the constriction A to the chamber outlet  52  is preferably 1 to 15 millimeters and more preferably 10 millimeters. The constriction width angle R 2  is preferably 0.00001 to 10 degrees and more preferably 0.0005 degrees. 
     Turning now to the relevant parameters of the piezoelectric body  32 , employing a thicker piezoelectric body  32  results in increased vibration which causes greater resonance to be produced thereby increasing fluid throughput. However, increased resonance may also result in increased negative pressure at the constriction A, thereby causing vaporization of dissolved gas. In order to increase the pumping rate, two or three piezoelectric bodies having thicknesses of 0.1 millimeter to 1 millimeter can be stacked. 
     Finally, with respect to the voltage V applied to the piezoelectric body  32  at the frequency n, increases in frequency and/or voltage with initially increase fluid throughput. Relative to the other optimized parameters herein provided for micropumps  30  according to the present invention, the frequency n is preferably in the range of 50 to 4500 hertz. With respect to voltage, the higher the voltage, the greater the amplitude at which the piezoelectric body will vibrate, which will cause an increase the resonance phenomena produced within the chamber, and increase fluid flow. However, strong resonance phenomena will also produce a greater negative pressure at the constriction A, which will result in vaporization of dissolved gasses. Accordingly, relative to the other optimized parameters herein provided for pumps according to the present invention, the voltage V is preferably in the range of 20 to 100 volts. 
     Furthermore, the manner in which the voltage is applied also influences the throughput of micropumps according to the present invention. For example, the negative pressure in the area of the constriction A and thus the vaporization of dissolved gas is controlled by controlling the rate of rise of the voltage V.  FIG. 4  shows the manner in which the voltage is preferably applied. In the graph provided in  FIG. 4 , the horizontal axis  60  represents the passage of time and the vertical axis  58  represents increasing voltage. Line  56  is provided as a reference line from which angle R 3  is determined. Relative to the other optimized parameters herein provided for pumps according to the present invention, in order to reduce or prevent vaporization of dissolved gas, angle R 3  is preferably in the range of 3 to 45 degrees. 
     The infusion pumps  10  according to the present invention are operable to provide fluid flow rates of up to approximately 280 milliliters per hour and greater. It will be understood, however, that the flow rate achieved by the infusion pump  10  is dependent upon the back pressure imparted by the patient&#39;s circulatory system. Accordingly, as the back pressure imparted on the infusion pump  10  increases, the control factors provided to the micropump  30  must be changed in order achieve the desired flow rate while overcoming such back pressure. For example, while maintaining all other design parameters and control factors constant, at back pressure of zero kilopascal, the infusion pump  10  according to the present invention may achieve a flow rate of approximately 780 microliters per minute at a frequency of 50 hertz, or Hz; 1,610 microliters per minute at a frequency of 100 Hz; and 1,930 microliters per minute at a frequency of 150 Hz. At back pressure of 60 kilopascal, the infusion pump  10  according to the present invention may achieve a flow rate of approximately 230 microliters per minute at a frequency of 50 Hz; 440 microliters per minute at a frequency of 100 Hz; and 630 microliters per minute at a frequency of 150 Hz. 
     In yet another embodiment of the present invention, as shown in  FIGS. 5   a  and  5   b , a micropump  330  comprises the piezoelectric body  32  attached to one side of the vibration layer  36 . An opposite side of the vibration layer  36  is affixed to the pump base  38 . The pump base  38  comprises the inlet channel  40 , the outlet channel  42 , and the chamber  44 . In contrast to the micropump  30  described above, the micropump  330  further comprises inlet valve  60  and outlet valve  62 . The inlet valve  60  and the outlet valve  62  are operable to transpose as shown in  FIGS. 5   a  and  5   b.    
     In operation, when the piezoelectric body  32  is manipulated in the direction of arrow  64 , chamber  44  becomes negatively pressurized and inlet valve  60  opens allowing fluid flow into the chamber  44 . Conversely, outlet valve  62  is pulled closed thereby discouraging any backflow of fluid in to the chamber  44 . When the piezoelectric body  32  is manipulated in the direction of arrow  66 , chamber  44  becomes positively pressurized. Inlet valve  60  is pushed closed thereby preventing flow out through the inlet channel  40 . Conversely, outlet valve  62  is opened to allow fluid flow through the outlet channel  42 . 
     It will be understood that features, such as the valves  60  and  62 , described above with respect to micropump  330  may be combined with any of the design features described with respect to micropumps  30  and vice versa in order to achieve micropumps according to the present invention. 
     As shown in  FIG. 6 , the infusion pump  10  according to the present invention further comprises a first circuit  70 , a second circuit  72  and flow meter  74 . The first circuit provides power to the electrode  34 , shown in  FIG. 2   a , and the second circuit  72  provides power to the flow meter  74  and AC/DC conversion of the data from the sensors associated with the flow meter  74 . The first and second circuits  72  and  74  are accordingly in electrical communication with a power source such as a wall mounted plug, a battery, or a combination thereof. The obvious advantage of an infusion pump  10  employing a battery is the corresponding mobility of the infusion pump  10  while maintaining uninterrupted flow of infusion fluids. 
     The flow meter  74  associated with the infusion pump  10  may comprise a variety of know flow meters. The flow meter  74  may comprise a variety of known flow meters. For example, the flow meter  74  may be configured to determine fluid flow rates by employing a heater that heats the fluid being monitored and senses the flow of the heated fluid downstream of the heater. Such flow meters are available from Sensirion AG of Switzerland and Siargo Incorporated of the United States of America and are described in greater detail in at least U.S. Pat. No. 6,813,944 to Mayer et al. and U.S. Publication No. 2009/0164163, which are herein incorporated by reference. Alternatively, the flow meter  74  may be configured to employ two pressure sensors positioned on each side of a constriction within the fluid flow path. Fluid flow rates are determined by the relative difference between the pressure sensors and changes thereof. 
     It is noted that while infusion pump  10  has been depicted as a single unit or component in  FIGS. 1 and 6 , certain elements of the infusion pump  10  may be compartmentalized within different components or bodies. For example,  FIG. 7  shows an infusion pump  10  in which first and second circuits  70  and  72 , respectively, housed in body  20   b  separate from the micropump  30 ,  330  and the flow meter  74  which are contained in body  20   a . Such as system is advantageous for several reasons. For example, it may be preferable to fabricate body  20   a  comprising the micropump  30 ,  330  and the flow meter to be disposable. Accordingly, risk of contamination and the costs associated with cleaning and preparing the micropump  30 ,  330  and flow meter  74  for use with different patients can be decreased. However, in order that the cost of production of the infusion pump may be minimized, it may be desirable to make as few as possible of the components of the infusion pump  10  disposable. Accordingly, body  20   b  in which the circuits  70  and  72  are contained may be considered non-disposable, or reusable. 
     The infusion pump  10  of the present invention comprising multiple bodies, such as bodies  20   a  and  20   b , may employ structures to establish electrical communication between one another. For example, as shown in  FIG. 1 , body  20   a  may comprise a communication port  26  through which the first and second circuits  70  and  72  may be placed in electrical communication with the electrode  34  of the micropump  30  and the flow meter  74 , respectively. The communication port  26  may be fabricated from a variety of electrical connectors/terminals and wire types known in the art. 
     Alternatively, as shown in  FIG. 8   a , the body  20   a  and the body  20   b  may be formed such that the bodies can be physically attached to one another. For example, body  20   a  comprises recesses  76  formed substantially around a perimeter of the body  20   a . Body  20   b , in turn, comprises binders  78  positioned also substantially around a perimeter of the body  20   b . The binders  78  are deflectable and shaped to fit into the recesses  76  formed in the body  20   a . When the disposable body  20   a  is to be mated with the non-disposable body  20   b , the two bodies are pressed against one another thereby forcing the binders  78  to deflect outward around the base  84  of the body  20   a  until the binders  78  engage the recesses  76  of the body  20   a . Once the binders  78  of the body  20   b  engage the recesses  76  of the body  20   a , the two bodies are pulled towards one another and maintained in friction fit. 
     Electrical communication is established between bodies  20   a  and  20   b  through complementary electrodes  82  formed on the surface  86  of the body  20   a  and the surface  88  of the body  20   b.    
     In certain embodiments of the present invention, is may be desirable to seal the electrodes  82  within a fluid tight environment. Accordingly, seal  80  may be embedded in the surface  86  of the body  20   a  and/or in the surface  88  of the body  20   b . Upon mating of the two bodies  20   a  and  20   b , the seal  80  is engaged and a fluid tight environment is established around the electrodes  82 . The seal  80  may comprise a soft, yielding material such as Teflon, silicone, or other polymeric material. In order to prevent the seal  80  from moving, the seal may be embedded at least halfway into the surface in which it is attached. 
     It will be noted that because the piezoelectric body  32  vibrates during operation, the construction and mating of the bodies  20   a  and  20   b  must be sufficient to maintain the electrical communication between the bodies. For this, it is necessary that any wires with which the piezoelectric body  32  is connected to the electrodes of the body  20   a  be made from a material that is flexible, and which bonds strongly with the piezoelectric body. 
     In certain embodiments of the present invention, a user interface or indicator, not shown, is provided on the body  20   b . The user interface may provide information such as the name and specifications of the fluid or pharmaceutical(s) being infused by the pump. The user interface may also be color-coded, so that the type of drug can be easily recognized. If multiple drugs are to be injected, in a system in which multiple infusion pumps  10  are employed, the user interface may provide pump identifiers that allow the user to easily identify a specific pump. The user interface may also provide warning lights that alert a user of particular operation parameters or operating conditions that have been encountered. 
     In yet another embodiment of the present invention, infusion pump  10  further comprises a control terminal, not shown, that is in electrical communication with the first and second circuits  70  and  72 . The control terminal serves to provide control factors to the micropump  30 . Micropump  30  control factors include, for example, the voltage V applied to the piezoelectric body  32  at the frequency n and the manner or rate at which the voltage V is increased and decreased during operation. The control terminal is in wireless or wired communication with the first and second circuits  70  and  72 . In a one embodiment, a nurse or other caregiver may employ one control terminal to independently establish electrical communication with and thereby control a single infusion pump  10 . Once the control factors have been provided to the infusion pump  10 , the same control terminal may be carried or otherwise transported to a different infusion pump  10  in order to provide the control factors for the second infusion pump  10 . Stated alternatively, it is contemplated that, for example, in a hospital setting, a plurality of different infusion pumps  10  can be controlled by a single control terminal. In such an embodiment, the control terminal may also serve to receive fluid flow rate data from the flow meter  74  associated with the infusion pump  10 . 
     In another embodiment, the control terminal comprises a centralized patient fluid management system that is associated with a single patient during the course of the patient&#39;s treatment, i.e. the control terminal is not shared between multiple patients. In such an embodiment, the control terminal is operable to provide control factors to a plurality of infusion pumps  10 , receive flow rate data from the flow meter  74 , receive patient biological data from various patient sensors associated with the control terminal, and to provide various patient information to the caregiver based upon such patient data received. Such a control terminal is described in greater detail in the Assignee&#39;s U.S. patent application Ser. No. 12/972,374, entitled Patient Fluid Management System, filed Dec. 17, 2010. 
     Of particular significance, is the fact that the micropump  30  of the infusion pump  10  does not exchange data with the control terminal, i.e. the electrical communication between the micropump  30  and the control terminal is one-way, from the control terminal to the micropump  30 . The micropump  30  is a slave to the control terminal. This configuration is advantageous because it provides for a simplified and more economical infusion pump  10 . For example, by making the micropump  30  a slave of the control terminal, the circuitry within the micropump  30  is simplified and thereby more economical to manufacture. In view of the above described embodiments in which the micropump  30  is disposable, a hospital or clinic may more economically obtain the disposable portions of the infusion pump  10  and only have to acquire and maintain a limited number of the more complex and more costly control terminals. 
     Although the invention has been described in terms of particular embodiments and applications, one of ordinary skill in the art, in light of this teaching, can generate additional embodiments and modifications without departing from the spirit of or exceeding the scope of the claimed invention. Accordingly, it is to be understood that the drawings and descriptions herein are proffered by way of example to facilitate comprehension of the invention and should not be construed to limit the scope thereof.