Over one million surgical procedures in the United States each year involve bone and cartilage replacement (Langer et al.,  Science 920:260-266 (1993)).  The poor healing characteristics of cartilage have driven the demand for options to replace or supplement damaged tissue.  Synthetic materials have been employed for this purpose, but issues of biocompatibility and eventual material failure by fatigue or wear ultimately limit their use.  Allografts present an alternative, but their use can be limited by immunological complications, transmission of infectious diseases from the donor, premature resorption of the transplant, and lack of the availability of donor material.  As a result, the use of autologous cartilage and/or bone grafts is considered a primary option (Lovice et al.,  Otolaryngol.  Clin.  N. Am.  32:113-139 (1999)).  This approach is hindered by the clinical difficulties associated with harvesting donor tissue.  Tissue engineering has been proposed as an alternative route by which tissues are regenerated by cells that are seeded into biodegradable polymer scaffolds that present an appropriate chemical and physical environment for the tissue growth either in vitro or in vivo after re-implantation. 
Each year, millions of people suffer severe acute or chronic cutaneous wounds.  While great progress has been made in both the fundamental understanding of the biology of wound healing and the clinical treatment of wounds, there are large margins for improvement: acute wounds still require many weeks of treatment, and chronic wounds associated with old age and diabetes still often persist indefinitely.  In both the study and treatment of wounds, scientists and doctors lack tools with which to manipulate the wound environment with high spatial and temporal precision.  In order to fully control the wound healing process, one must be able to deliver and extract reagents with micrometer-scale spatial resolution (the scale of individual cells) over the macroscopic dimensions of a typical wound, and with minute to hour-scale temporal resolution over the days to weeks of healing. 
Tissue engineering holds promise as an approach to generate replacement tissues and organs for those lost by injury or disease.  Particular progress has been made in musculoskeletal tissues such as cartilage (Brittberg, M., Clinical Orthopaedics and Related Research 367 Suppl.:S147-155 (1999)) and bone (Vacanti et al.,  New Eng.  J. Med.  344:1511-1514 (2001)); for these systems, limited clinical success has been achieved.  Nonetheless, engineering tissue has been hindered by the lack of sophisticated tools for tailoring the physical and chemical environment of the tissue-forming cells.  Recent work has demonstrated success in growing cartilage in 3D scaffolds with physiologically appropriate size and shape; this process is based on injection molding of chondrocyte-seeded gels (Chang et al.,  J. Biomed.  Mat.  Res.  55:503-511 (2001); Chang et al.,  Plastic and Reconstructive Surgery 112:793-799 (2003)).  Recent advances have also been made in the field of microfluidics, allowing for the control of fluids on micrometer-scales within organic materials (Stone et al.,  Annual Review of Fluid Mechanics 36:381-411 (2004); McDonald et al.,  Anal.  Chem.  74:1537-1545 (2002)). 
The goal of tissue engineering is to initiate and direct the growth of living tissue for applications which include: studying of basic biological questions, in vitro testing of drugs and environmental agents, and, ultimately, replacing the form and function of compromised tissue in the body by surgical transplantation.  Enormous progress has been made over the past few decades toward this goal, with some engineered tissues having entered the clinic (Langer et al.,  Principles of Tissue Engineering, ed.,  Academic Press: San Diego (2000).  A central aspect of successful strategies in tissue engineering is the preparation of an appropriate chemical and mechanical environment in which to grow the tissue cells.  Ideally, this environment should be able to mimic aspects of the native environment in which the tissue of interest would have developed in vivo.  To this end, important work has been done to tailor the chemical character of the matrix in which tissue cells are embedded (Rowley et al.,  Biomat.  20(1):45-53 (1999)), to assess the effects of physical stimuli (e.g., mechanical and electrical) (Bonassar et al.,  J. Ortho.  Res.  19(1):11-17 (2001)), and to assess the effects of soluble chemical stimuli such as growth factors in the media surrounding the growing tissue (Sweigart et al.,  Tissue Eng.  7(2):111-129 (2001)). 
In attempting to tailor the environment of a developing tissue, serious consideration must be made of mass transfer to and from each cell in the system.  The chemical input and output of a cell is crucial for its basic metabolic functions, and for its interactions with the outside and its neighboring cells.  From an engineering stand point, the transfer of soluble species to and from cells in the tissue is one of the principal ways in which information can be delivered and extracted from the tissue, in order to influence and monitor its development.  While questions of mass transfer are often discussed in tissue engineering context, there is a lack both of tools with which to implement controlled mass transfer in a growing tissue, and of basic design rules for such a control system. 
The creation of a synthetic scaffold that provides the appropriate structural and chemical environment to developing cells is a core strength and challenge of the tissue engineering approach.  A variety of synthetic and naturally occurring polymer scaffolds have been used to define both the macroscopic shape and chemistry of the solid structure in which cells can bind (Frenkel et al.,  Ann.  Biomed.  Eng.  32:26-34 (2004)).  Many materials require that cells be seeded into a preformed structure.  This post-seeding method has a distinct disadvantage in that the seeding density is typically inhomogeneous, at least initially (Obradovic et al.,  Aiche Journal 46:1860-1871 (2000)).  An alternative approach has been introduced based on polymers such as alginate (Chang et al.,  J. Biomed.  Mat.  Res.  55:503-511 (2001) and agarose (Hung et al.,  J. Biomech.  36:1853-1864 (2003)) that can be solidified under physiological conditions; this method permits the cells to be uniformly suspended in the polymer prior to gelation, leading to a highly homogeneous initial distribution of cells within the scaffold.  The form of the polymer-cell gel can be imposed by casting or molding (Chang et al.,  J. Biomed.  Mat.  Res.  55:503-511 (2001); Chang et al.,  Plastic and Reconstructive Surgery 112:793-799 (2003)).  The use of alginate for long term culture of chondrocytes in vitro has been documented extensively (Beekman et al.,  Exp.  Cell Res.  237:135-141 (1997); Guo et al.,  Connect Tissue Res.  19:277-297 (1989)).  In vivo, alginate has been successfully employed as an injectable vehicle for chondrocyte delivery in the treatment of vesicoureteral reflux (Atala et al.,  J. Urology 150:745-747 (1993)). 
A common strategy for controlling the chemical environment of a tissue scaffold is to implant it in a living animal such that the animal's body supplies the basic nutrients, and, perhaps appropriate signals to encourage development (Chang et al.,  J. Biomed.  Mat.  Res.  55:503-511 (2001)).  This technique can pose challenges due to immunological rejection, resorption, and inaccessibility for detailed study of tissue development.  As an alternative, in vitro bioreactors have been designed to control the physical and chemical environment of the developing scaffold (Martin et al.,  Trends in Biotechnology 22:80-86 (2004)).  In the engineering of cartilage, fluid motion (Freed et al.,  Journal of Cellular Biochemistry 51:257-264 (1993)) and mechanical deformation (Bonassar et al.,  J. Ortho.  Res.  19:11-17 (2001)) have been explored as functional characteristics of bioreactors.  Fluid motion is introduced primarily to assist mass transfer from the culture medium into scaffolds, but hydrodynamic stresses appear to influence the development of tissue as well (Martin et al.,  Biorheology 37:141-147 (2000)).  A variety of modes of fluid motion have been explored: spinner flasks, rotating wall vessels (Martin et al.,  Trends in Biotechnology 22:80-86 (2004)), and perfusion reactors (Pazzano et al.,  Biotechnology Progress 16:893-896 (2000)).  No technique has yet been presented that allows for fluid motion to be directed along well-defined paths within a material suitable for use as a tissue scaffold. 
The development of microtechnology for the control of fluid behavior has been focused over the past decade on developing small-scale chemical systems for analytical and synthetic manipulations (Stone et al.,  Annual Review of Fluid Mechanics 36:381-411 (2004); Whitesides et al.,  Phys.  Today 54:42-48 (2001)).  Recently, several groups have also developed microfluidic systems for controlling the culture environment of one or few cells grown in monolayers within the microchannel (Takayama et al.,  Proc.  Nat'l Acad.  Sci.  U.S.A. 96:5545-5548 (1999)).  A network of microfluidic channels in a silicon wafer also been used as a substrate on which a monolayer of capillary endothelial cells were grown (Borenstein et al.,  Biomedical Microdevices 4:167-175 (2002)).  The group of Borenstein has recently demonstrated the fabrication of microchannels in poly(DL-lactic-co-glycolide) (PLGA) (King et al.,  Advanced Materials 16:2007-2012 (2004)); these microfluidic structures may be appropriate for use as scaffolds for 3D culture of cells.  In general, there is an outstanding challenge to apply microfluidic methods to control the chemical environment of cells in a 3D culture. 
The present invention is directed to overcoming these and other deficiencies in the art.