Fluid flow control is an essential part of medical devices such as intravenous infusion pumps and enteral feeding systems. These fluid flow control systems must meet a complex and conflicting set of requirements, such as broad flow rate range, wide ranging fluid viscosity, inevitable presence of harmful amounts of gas, changing source pressure, changing patient pressure, variable patient line resistance, and a wide range of tubing configurations.
Reliability and ability to detect fault conditions are critical features of such flow control devices. Low acquisition and maintenance costs are important characteristics also.
The usability of the system is vitally important, as it impacts the workflow of caregivers, which has a strong, but indirect, impact on the quality of patient care. This usability includes ease of loading the sterile tubing set, the need for attention from the caregiver during the fluid delivery period, and attending to unnecessary alarm conditions.
Conventional fluid control or pumping mechanisms suffer from an unfavorable tradeoff between sophistication and complexity. The added complexity of many modern systems has led to a lack of reliability, resulting in product performance failures, high levels of maintenance, product recalls by regulatory agencies, and documented high rates of patient harm.
One of the earlier types of fluid pump, as marketed by Harvard Apparatus Company and as replicated in the market hundreds of times thereafter, is a syringe pump. In a syringe pump, fluid is contained within a commonly found glass or plastic syringe, manufactured with a well-specified diameter and stroke length. These are the same syringes that are used to provide manual injections of sterile fluid. The piston of the syringe is securely held and, usually with a lead screw mechanism, the piston is advanced in carefully timed steps of a motor. Each step of the motor expresses a known amount of liquid out of the syringe and into a line going to the vasculature of the patient. The syringe pump offers a very simple mechanism and an extraordinarily simple control system, consisting of a timer circuit, set by the desired fluid flow rate. Force and position sensors are often added to provide feedback regarding occlusions, misloading, and end of infusion. The syringe pump design is inherently limited, however, by the relatively small size of the syringe, in the amount of fluid infused and in the maximum fluid flow rate, so this design does not satisfy the needs of many clinical applications. Ironically, at the very small volumes and flow rates, the syringe pump suffers from a discontinuity of fluid flow, based on the high static friction of the syringe. Very small movements of the drive motor do not necessarily translate into movement of the piston and delivery of fluid; it may take multiple motor steps and multiple time intervals before the piston actually delivers fluid to the patient. Long delay periods between delivery are not desirable clinically. A further deficiency in the syringe pump is the improper impedance match with the patient's vasculature; the syringe pump motor drive is equipped with a motor that is capable of reliably meeting the maximum torque foreseen by the system. This powerful motor is also geared down such that very low displacements can be achieved, giving the pump the ability to deliver at low flow rates. The combination of the powerful motor and the gearing, however, allows the syringe drive to generate fluid pressures that are far in excess of those needed to safely infuse a fluid into the vasculature of a patient. The consequence of this potentially high pressure output is that harmful levels of fluid pressure can be applied to the patient, with deleterious effects, especially in the event of an extravasation of the infusion catheter or the creation of a bolus upon release of a temporary occlusion.
Variations of the syringe pump are to be found in the form of a reciprocating piston that can draw from a fluid bag or vented bottle. Such devices, as found with the Abbott/Hospira Plum™ infusion device, overcome the volume limitation of a syringe pump. Added complexity for valving serves to increase cost and reduce reliability. A large volume pump, because of its multiple fluid connections and air spaces, creates an environment, not found with syringe pumps, for the introduction of harmful air bubbles, which must be detected and accommodated. These reciprocating piston pumps still retain the disadvantage of impedance mismatch described above for syringe pumps.
The most common form of infusion pump is the peristaltic pump, whereupon fingers or rollers occlude a section of flexible tubing in peristaltic fashion, expressing fluid out the tube toward the patient. This mechanism provides the simplest configuration to carry the sterile fluid in the form a simple flexible tube. The peristaltic pump suffers the same impedance mismatch fate as the syringe pump, because the forces required to faithfully occlude a portion of the flexible tube are great, allowing the pump to generate harmfully high infusion pressures. This potentially high pressure can be mitigated through the use of force sensors on the tubing, adding complexity and cost. The problem with air ingress to the patient is the same as with the reciprocating piston pump described above. The peristaltic pump introduced a new problem related to fluid flow accuracy, since the amount of fluid expressed to the patient is entirely dependent on the interior diameter of the fluid tubing in its uncompressed state. In fact the surface area error is a square law function of the error in the diameter, so a 10% error in the diameter would yield an unacceptable 21% (1.12) error in the volume expressed to the patient. Unfortunately, there are two very common events that can reduce the effective diameter of the tubing: one is the fatigue of the tubing as it is repeatedly worked by the peristaltic mechanism and the other is the failure of the tubing to refill completely due to low flow from the fluid source.
There is another class of pumps providing single flow rates using a constant force spring, membrane, or gas reaction pushing fluid against a fixed, calibrated resistance. These devices do not provide the programmable variation of flow rate needed for most clinical applications.
One variation of the reciprocal piston pump was designed and marketed by FluidSense Corporation of Newburyport, Mass. It used a flexible membrane connected to a spring-loaded piston on one side and sterile fluid on the other. A low cracking pressure passive inlet valve and an actively operated momentary outlet valve provided for a pumping action if the spring loaded piston were “cocked” back to load the spring, providing a positive fluid force. A highly sensitive linear encoder was used to watch the position of the spring-loaded piston, providing information on the fluid pressure and volume. This design allowed for a simplified and more sensitive pump mechanism, but the flow was intermittent with the action of each pulse of the outlet valve and the driving pressure varied from 3 to 7 PSIg, higher than necessary for most clinical applications. It also suffered from the introduction of air bubbles, as with all large volume pumping systems.
Programmable infusion devices, as opposed to single rate delivery systems, all suffer from two effects of electromechanical complexity. First, there are usually tight mechanical tolerances which can be disturbed by shock, vibration, temperature shifts, and aging. Infusion pumps are often out of their performance specifications, sometimes intermittently, making troubleshooting very expensive and difficult. Secondly, these complex mechanisms are often difficult to disinfect. Customers have only recently become sensitized to the extremely high importance of disinfecting infusion pumps and other medical devices. Cross contamination of patients is one of the top healthcare issues in the acute care environment.
Another particular problem that patients and caregivers face with great regularly is the presence of air bubbles in the fluid path. Conventional infusion pumps observe a segment of tubing via an ultrasonic or optical detector circuit. They reliably detect bubbles with high sensitivity. Unfortunately, the specificity of these sensors is low, so false alarms are commonplace. When these bubbles are detected, three bad things happen. First, the pump goes into an alarm condition and fluid flow to the patient is halted, which can often cause harm to the patient by withholding needed medication. Second, the alarm at the bedside causes significant distress to the patient and the patient's family. Third, the alarm disrupts the nurse's workflow, taking time away from other patients and directing the nurse's attention toward the infusion pump and away from the patient.
Air eliminating filters are commonly found in infusion therapy administration sets. These filters fail to solve the problems identified above, because these filters do not function properly when exposed to negative gauge pressures if they are positioned proximal to the infusion pump. If these filters are placed below the infusion pump, then there is no way for the pump to verify that these filters are in place, so the alarms must still stay active. These filters must also incorporate hydrophilic filters, which are not compatible with certain medical fluids, such as whole blood.