Existing drug-delivery technology is typically not appropriate for long-term programmable infusion into certain body parts, such as, for example, the inner ear. Known drug-delivery devices include external and implanted infusers, osmotic pumps, and erodible polymer-drug systems. These systems range from passive devices, which have a low level of predictability in their dispensing rates, to electronically-controlled rate dispensers, and finally to fully programmable infusers.
Device volumes range from approximately “pill” sized to over ten cubic inches, generally depending on their maximum dispensing volume and the sophistication of their control systems. Though small in volume, erodible polymer and porous membrane systems must typically be utilized to deliver a specific compound or, at best, a set of compounds with similar chemistry and transport properties. These devices are generally short to medium term delivery devices (i.e., less than approximately six weeks) with unalterable, non-constant delivery profiles. Existing osmotic pump-based delivery systems are similar in terms of device size and lifetime, and they too are typically only capable of fixed rate delivery. Other available devices may offer more sophisticated control, and may be effective for the treatment of disorders such as chronic pain, but these systems generally use macro-scale, conventionally fabricated pumps that are relatively large. As a result, they are generally only practical when implanted in subcutaneous tissue in the torso.
In some macro-scale systems (for example, pressure-regulation systems), bellows are used with fast valves to damp out rapid changes in pressure. Miniaturized bellows have been fabricated, but reliability and cost make them inappropriate for some applications. Other systems have used compliant tubing as a capacitive element. In such systems, as pressure increases, the tubing volume generally expands by an amount dependent on the tubing material and its dimensions. However, achieving appreciably large capacitance values using standard tubing requires a relatively large total tubing volume. Moreover, the tubing typically cannot have sufficiently thin walls to offer the values of capacitance desired within a small volume. Thus, relatively long and thick tubing sections are required, thereby increasing the total system volume of the injection system.
Small size is often important for many applications, particularly in implantable drug-delivery systems where both the trauma from the surgical implantation and foreign body response are reduced when the implant size is reduced. In addition, large tubing surface areas allow for significant vapor or gas permeation, which is detrimental to a system with low fluid volumes, particularly if they are not refilled. Integration of tubing as a capacitive element in a microfluidic system also requires connectors from the tubing to the other elements (e.g., pumps, valves, etc.). These connectors add to system volume and are a source of both leakage and dead volume. Yet another disadvantage in known microfluidic systems is their sensitivity to the tubing dimensions and mechanical properties, which may change over time.
In addition, a MEMS-based drug-delivery device may require a valve system for fluid control within the device. Exemplary valve systems known in the art are depicted in FIGS. 1A-2B. For example, FIGS. 1A and 1B depict a valve system 100 that includes a pair of four-way stopcock valves 120, 140. In this exemplary system 100, during filling of a fluidic manifold or loop 110, as depicted in FIG. 1A, one valve 120 typically connects a load reservoir (not shown) to the manifold 110 through a load reservoir connector 130, while another valve 140 vents the other end of the manifold 110 though a venting connector 150. As depicted in FIG. 1B, during manifold circulation in the direction indicated by the arrows 160, the two stopcock valves 120, 140 are rotated to join the input and output sections of the manifold or loop 110.
FIGS. 2A and 2B depict an exemplary rotary-type valve system 200. During loading, as depicted in FIG. 2A, a load reservoir connector 130 and an inlet 210 of the manifold 110 are connected through a valve 230, and an outlet 220 of the manifold 110 is vented through a venting connector 150. During circulation, as depicted in FIG. 2B, the manifold inlet 210 and outlet 220 may be connected to valve chambers within the rotary valve 230 and to a connector 240, external to the valve 230, which joins those valve chambers.
Unfortunately, the size of both the stopcock valves 120, 140 and rotary valves 230 is large (typically greater than one inch for the stopcock valves 120, 140 and greater than two inches in the maximum dimension for the rotary valves 230) relative to the overall desired system size. In particular, their size is large for typical implantable drug-delivery systems, and also for drug-delivery systems mounted on small lab animals typically used for preliminary animal studies. In addition to these large physical dimensions, stopcock valves have relatively large internal fluid volumes (e.g., 100 microliters or greater, including fitting volumes), which is generally undesirable when the total fluid volume of the system is desired to be small. Typical Luer fittings for stopcock valves are designed to mate to similarly sized (i.e., not microscale) fluidic components, and thus also add considerably to system size and internal fluid volume.
As a result, there is a need for both a capacitive element and/or a valve system that is small enough to fit within a microfluidic system, such as an implantable drug-delivery system, while still being capable of performing fluid control functions with sufficient reliability and efficiency.