Multi-port fluidic wound dressing device

Abstract

A fluidic wound dressing device comprises a branching network of fluidic channels connecting a single inlet port to a plurality of outlet ports in a biocompatible, treatment site-facing surface. The branching network enables the delivery of injected fluids containing therapeutic agents such as cells, drugs, antimicrobics, saline, and the like to the treatment site vis the outlet ports with a substantially uniform distribution. The outlet ports may be arranged in a regular grid array or other arrangement of substantially consistent density. The branching network may be based on a fractal arrangement such as an H-network to provide channels of substantially equal length connecting the inlet with each outlet. The relative cross-sectional areas of the channels may be determined in accordance with Murray's law to provide for optimal fluid delivery. The device may be used to deposit hydrogels comprising therapeutic cells such as fibroblasts and keratinocytes in layered structures.

Description

MULTI-PORT FLUIDIC WOUND DRESSING DEVICECross-Reference to Related Applications

[0001] This application claims priority to and the benefit of United States Provisional Application No. 63 / 580,578 filed September 5, 2023, the entirety of which is incorporated herein by reference.Technical Field

[0002] The present disclosure relates to a wound dressing device, and in particular to a fluidic dressing device with multiple ports for delivering therapeutic agents to a treatment site as in situ bioprinting.Technical Background

[0003] As the largest human organ, the loss of skin integrity from an illness or injury causes a significant physiologic imbalance that can lead to serious impairment or even death.Extensive and deep wounds often require the transplantation of autografts or allografts for closer, but these methods are limited by donor tissue shortages, infection risks, and immunogenic rejection. A promising alternative is the fabrication of tissue engineering scaffolds (engineered skin grafts), particularly for large, deep, and / or chronic wounds. However, fabricating scaffolds in custom sizes and dimensions to adequately cover wounds with varying depths or topographies is challenging. Additionally, the time required to generate enough cells can be extensive (1-3 weeks). Finally, the insufficient availability of nutrients and oxygen, especially in the central regions of large scaffolds, compromises their overall therapeutic efficacy, which may lead to poor regenerative outcomes.

[0004] Conventional three-dimensional (3D) bioprinting is a fabrication technology that offers high control over the spatial arrangement of biomaterials and cells, which may be used to facilitate the preparation of skin substitutes that mimic the properties of host tissue and facilitate in situ skin regeneration. However, current bioprinting systems are often limited for point-of-care use. Typically, such systems involve the ex vivo fabrication of the skin graft, and requiring lengthy periods before application to the wound. This “print-then-transplant” approach has several drawbacks. The extensive time needed to print means the patient can experience scarring or infection while waiting, as well as potential problems of delayed wound closure, increased clinical protocol complexity, extended hospital stays, and additional costs. Further, manual handling of the fragile printed graft duringtransfer to the wound site limits material composition and increases the risk of contamination and infection. Additionally, tissue regeneration may be hindered because external bioreactors used during maturation do not entirely mimic the native physiological environment.Brief description of the drawings

[0005] In drawings which illustrate by way of example only embodiments of the present invention,

[0006] FIG. 1 is an exploded view of an example multi-port fluidic wound dressing device with four outlet ports.

[0007] FIG. 2 illustrates a set of H-channel network designs employed in fabricated devices, ranging from 4 to 128 outlet ports.

[0008] FIG. 3 is a diagram of one portion of an H-channel network.

[0009] FIG. 4 is a set of images and a graph depicting results of cell delivery through an example 4-port device.

[0010] FIG. 5 is a set of schematics, images, and a graph depicting in vitro dispersal of cells using the device.

[0011] FIG. 6 is a set of images of example multi-port fluidic wound dressing devices with increasing numbers of outlet ports.

[0012] FIG. 7 is a set of images and a graph depicting live cell count and proliferation after delivery of cells through an example 8-port device.

[0013] FIG. 8 is a set of images depicting results of deposition of cells using an example 32- port device.

[0014] FIG. 10 is a set of images depicting results of repeated deposition of layers containing cells using an example 8-port device, compared to a control and 3D bioprinter.

[0015] FIG. 11 is a set of images depicting further results of repeated deposition of layers containing cells using the example device, compared to a control and 3D bioprinter.

[0016] FIG. 12 is a set of images depicting the result of co-cultured deposed layers of fibroblasts and keratinocytes by an example device, compared to a control.

[0017] FIG. 13 is a set of images depicting an example device with drainage holes and a complementary absorbent pad.Detailed description of example embodiments

[0018] The present disclosure accordingly provides a novel multi-port, fluidic wound dressing device, which contains fluidic networks connecting an inlet port with multiple outlet ports to deliver fluids and therapeutic agents (e.g., hydrogels, cells, scaffolding, growth and / or differentiation factors, antimicrobial agents, drugs, saline flush) in situ directly to a treatment site (e.g., a wound surface). The device can be used to deliver the fluids substantially uniformly over a large treatment area, through the use of a branching network of channels that provide substantially uniform flow rates to each outlet port. As such, the device may enable early delivery of therapeutic agents such as autologous cells (e.g., on day 1 ) to improve treatment outcomes. The device may further be used to deposit multiple injections of therapeutic agents in a continuous or semi-continuous manner, or to fabricate single or multi-layered skin structures to facilitate wound healing or other tissue engineering applications, potentially augmenting wound healing rates and mitigating scar tissue formation and enabling optimized treatment schedules accordingto the nature and size of the wound. Further, as the wound dressing itself comprises a device for distribution of therapeutic agents while maintaining a protected, closed environment at the treatment site, it may reduce infection risk.

[0019] An example device 100 is depicted in exploded view in FIG. 1 . In one fabrication, the device 100 is manufactured in multiple layers defining the inlet port, outlet ports, andconnecting network of fluidic channels. These layers may be approximately 1 mm in thickness. A first layer 10 comprises a plurality of outlet ports 15, which may simply be apertures punched or otherwise formed in the first layer. In the example of FIG. 1 , only four outlet ports 15 are provided arranged in a substantially regular grid; but as discussed below, significantly more outlet ports 15 may be provided. The outlet ports 15 are distributed on the first layer 10, defining an approximate treatment area on the exterior surface of the first layer 10. Also as will be discussed below, the treatment area can be determined by the spacing and the number of outlet ports 15, and the device 100 is therefore scalable to any suitable size. A second, intermediate layer 20 comprises a network 25 of branching fluidic channels, the terminals of which generally correspond to the locations of the outlet ports 15 in the first layer. A third layer 30 comprises a single inlet port 35, generally corresponding to a central location of the treatment area defined by the outlet ports 15 and coincident with a principal channel 27 of the network 25. The inlet port 35 may comprise a simple aperture, optionally with reinforcement or attachment means for receiving a syringe, tube, or other equipment used to introduce fluids and therapeutic agents into the device 100. When assembled, the inlet port 35 is in fluid communication with each of the outlet ports 15, and the introduced fluids and agents can then flow through the branched channels of the second layer to exit through a corresponding outlet port 15. While the device 100 in FIG. 1 is illustrated in three distinct layers, those skilled in the art will appreciate that the device 100 may be manufactured in different ways, in layers or otherwise, to provide a network of branching channels in fluid communication with the inlet and outlet ports. For example, as described below, an example device was fabricated in two layers, with the outlet port layer being molded with the channels. Alternatively, the third layer comprisingthe inlet port may be molded integrally with the channels.

[0020] At least the first layer, and particularly the surface intended to contact the treatment site, is biocompatible, that is to say, formed of a material suitable for application to a treatment site such as a patient’s skin surface or a wound site. Suitable materials include biocompatible polymers such as polydimethylsiloxane (PDMS) silicone, lubricant-infusedpolydimethylsiloxane (iPDMS), or polyurethane. The remaining body of the device 100 may be formed of the same material.

[0021] To facilitate substantially uniform distribution of fluid form a single injection inlet port over a relatively large area (e.g., up to 200 cm2or even larger), each branching channel between the inlet port 35 and an outlet port 15 is of substantially the same total length and volume as the others, such that fluid entering the device can exit through the outlet ports 15 at a similar volumetric flow rate. At the same time, the outlet ports 15 are distributed throughout the treatment area at a substantially consistent density for effective treatment. To achieve this over a large area — for example, in a 10 x 10 cm2area — the outlet ports 15 are distributed in several rows and columns. As a result, some outlet ports 15 are closer to the inlet port 35 than others, when the direct distance between their centers is measured. The network 25 of branching channels ensures that the overall channel lengths for each individual outlet port 15 is substantially the same. To achieve this, the network is designed as an “H-channel” (or H tree or H fractal) network, in which an initial channel branches into two perpendicular channels, which in turn may branch into a further two perpendicular channels, and so on. This can be appreciated more fully in FIG. 2, which schematically illustrates higher-order H-channel networks 25 and corresponding outlet port 15 positions at the terminals of each branch of the network 25. Example (a) is a first-order network, similar to that shown in FIG. 1 , with four terminals corresponding to four outlet ports 15. Examples (b) through (e) depict second, third, fourth, and fifth-order networks terminating at eight, sixteen, thirty-two, and sixty-four outlet ports 15. Thus, a complete H-channel network may support up to 2n+1outlet ports 15 for a single given inlet port 15, where order n corresponds to the number of nodes or junctions between the center of the network (corresponding to the location of the inlet port 35) and the terminal of a given branch (corresponding to the location of an outlet port 15). FIG. 3, which is a schematic of half of the second-order network shown in FIG. 2(b), demonstrates that there are two nodes (n1 and n2) between the network center (corresponding to the inlet port 35) and the terminal (corresponding to the outlet port 15). The rectangular nature of the H-channel network thus conveniently supports a regular grid array of outlet ports 15.

[0022] Additionally, the network 25 of branching channels may be sized to reflect Murray’s law in order to minimize resistance to flowthroughout the network. Murray’s law provides that in a network of tubular pipes in which a first branch (the parent) splits into N branches, flow resistance is minimized whenwhere rpis the radius of the parent branch and rdis the radius of a daughter branch. Murray’s law has been observed in nature, such as in the vascular and respiratory systems of animals, and has been demonstrated to provide maximum efficiency in the flow of liquids through branching networks. As the branching channels in the device 100 have rectangular cross-sections to facilitate fabrication, Murray’s law may be adapted as: 0.63where A is the cross-sectional area of the / th channel in the network, and there are only two daughter channels of equal rectangular cross-sectional area. Thus, for example, referring to FIG. 3, the ratio y of the cross-sectional area of channel D1 to the parent channel, and the ratio y of the cross-sectional area of channel D5 to D1 , are ideally approximately 0.63. Of course, the cross-sectional ratios of the parent and daughter channels in the network 25 need not follow Murray’s law, as suboptimal volumetric flow rate may still be acceptable for therapeutic treatments.

[0023] An example multi-port fluidic device 100 was fabricated using polydimethylsiloxane (PDMS) and a crosslinker (Sylgard™ 184). H-shaped channels were designed with different numbers of ports (orders n) and different cross-sectional aspect ratios (depth / width), then cut in a silicone sheet using a vinyl cutting machine to produce templates for forming channel networks with a depth of 0.5 mm or 1 mm. The PDMS and crosslinker, at a ratio of 10:1 , were fully mixed in a blender (FlackTek™ Speed Mixer™) for two minutes at 2000 rpm, then gently poured over the cut silicone sheet in a petri dish to completely cover the channel pattern. A plain PDMS layer was poured in a separate petri dish. After curing at60°C for 6 hours, the dried PDMS layers were peeled off the dishes and the inlet and outlet ports were punched through the layers (Miltex™ biopsy punch). The layers were cleansed using a basic oxygen plasma cleaning device (Harrick™ plasma cleaner PDC32G) for 60 seconds, introducing various chemical functional groups (silanol, SiOH) on the PDMS surfaces to be sealed together. Sealing the layers together produced the H-shaped channels embedded within the device and extending between the inlet port and outlet ports. The assembly was again cured, at 60°C for 12 hours.

[0024] Initially, 2-port and 4-port devices 100 (corresponding to 0thand 1storder networks) were fabricated with 2 mm wide initial channel widths, with daughter channel widths obtained from the above adaptation of Murray’s law, given that all channels had constant height. The inlet and outlet port diameters were 2 mm and 1 mm, respectively. The 4-port version of the device 100 had an area of about 8 cm2and an approximate 1 cm outlet port spacing. Individual branch segments (i.e., individual parent or daughter branches) were considered to have lengths L greater than the size of their junctions, such that the localized pressure losses at the junctions were insignificant in comparison to pressure drops over individual channels, and the total network volume V could be considered to be the sum of all individual channel volumes (A x L). Table 1 below sets out the geometric parameters of the 2- and 4-port devices 100:Table 1 . Geometric parameters of sample devices.

[0025] While the example devices 100 were manufactured with ports and channels in the millimeter range, depending on the intended application and fluids to be delivered by the device 100, these dimensions may be on the micron scale (e.g., ranging in the tens or hundreds of microns) or larger (e.g., around 10 mm). The uniformity of cell delivery andviability of cells delivered from the inlet port 15 to the outlet ports 15 was assessed using adult human dermal fibroblasts (HDFa, LONZA, Switzerland). Cells were cultured in growth medium (Gibco™ DMEM, 10% fetal bovine serum (FBS) and 1% antibiotic / antimycotic (Ab / Am) (Sigma Aldrich) until 80% confluency, and split into further passages by treatment with 0.05% trypsin-EDTA treatment (Sigma Aldrich). Cell passages 3-5 were used. HDFa at a concentration of 105cells / mL was injected into both the 2 and 4-port devices 100 at a flow rate of 5 mL / min for delivery onto culture plates coated with a 5 wt.% gelatin (Gel5) hydrogel. Control plates were those of the same cells and cell density, but delivered manually by pipette without the use of a device 100. All cells were incubated for 3 and 7 days, and stained using immunofluorescence techniques involving live / dead (Live / Dead® Cell Imaging Kit (488 / 570) (Life Technologies Inc.)) and DAPI (4',6-diamidino-2- phenylindole) staining. Generally, the same or similar methods were used throughout each investigation of device performance discussed below.

[0026] FIG. 4(a) shows the live / dead staining for cells delivered via an example 4-port device 100. Images (i), (iii), (iv), and (vi) labeled P1 , P2, P3 and P4 respectively depict the sample at each of the four outlet ports after 7 days. Image (ii) is a photograph of the sample device showing the locations of the four outlet ports. Image (v) depicts an area away from the outlet ports. Images (vii) and (viii) are 20X magnifications of the cells of outlet ports 3 and 4, respectively. FIG. 4(b) shows the result of the control sample at 5X and 20X magnification, while FIG. 4(c) shows the live and dead cell counts for each outlet port compared to the control sample. It was found that there was no significant difference in live cell density between the different ports. In the control sample, cells were observed to be randomly distributed, while in the device sample there was a much higher degree of clustering at the outlets, due to the localized delivery. Overall, the average cell density in the controls was about 1 .4 times lower than those at the outlet ports of the device, but cell viability exceeded 90% in both the control and device samples. In specific regions, cells elongated and formed connections, as indicated by white arrows in images (vii) and (viii) of FIG. 4(a), indicating cell proliferation.

[0027] In vitro cell dispersal was investigated by comparing human dermal fibroblasts (HDFs) delivered through the device 100 into wells to a control group of homogenous single cell delivery. A 5% wt. gelatin substrate layer was used to mimic a wound surface, and equal cell numbers were used for each cohort. The device was removed after 24 hours, and the petri dishes were labelled with DAPI and toluidine blue and imaged using confocal microscopy in order to count the cells. Fifteen locations, each with an approximate surface area of 4 mm2at the top, middle, and bottom sections of both the control and device samples were chosen for cell counting as depicted schematically in FIG. 5(a). In the device sample, some locations were located at device outlet ports, and others were located between ports. Cell counting employed systematic uniform random sampling to mitigate bias. FIG. 5(b) shows the average cell counts for the control (all locations), device (at ports), device (between ports), and device (average of all locations). FIG. 5(c) shows images of DAPI-stained cells (nuclei) for both the control and device (both between ports and at ports) at 5X magnification. It was found that the device samples had an average of 201 cells / mm2while the control sample had 100 cells / mm2within a 7-day incubation period. Furthermore, it was found that almost 95% of cells in the device sample were located around the outlets, with the remaining distributed around and underthe device, whereas in the control sample cells were distributed randomly everywhere with lower cell density in total after 7 days of incubation. This indicates that the higher cell density may enhance cell-to-cell interaction which leads to increased cell signaling leadingto this effect.

[0028] Additional devices 100 were fabricated with 8, 16, 32, 64, and 128 outlet ports to demonstrate the scalability of the device 100 to cover larger treatment sites, as set out in Table 2 below. The ratio of outlet to inlet port diameter was 2 (e.g., a 2 mm outlet port to a 1 mm inlet port). Photographs of these fabricated devices are provided in FIG. 6. Substantially uniform fluid delivery was verified by observing dyed liquid (food dye, ethanol, deionized water) injected into the device using a syringe pump.Table 2. Geometric parameters of sample devices.

[0029] HDFa were introduced into the 8-port device, using digital syringe pump at 10 mL / min, at a concentration of 106cells per milliliter. This was done on a substrate coated with a 5 wt.% gelatin hydrogel, and the cells were cultured for both 3 and 7 days. Live / dead staining and MTS (3-(4,5-dimethylthiazol-2-yl)-5-(3-carboxymethoxyphenyl)-2-(4- sulfophenyl)-2H-tetrazolium) proliferation assays were conducted and revealed that cell viability and proliferation rates were on par with control samples, which did not utilize the device. The absorbance recorded at490nm using a 96-well plate reader was indicative of the formazan product's quantity, as measured by absorbance at 490 nm, inversely correlated with the number of viable cells in the culture.

[0030] FIG. 7 shows Calcein AM stained living cells after delivery through the 8-port device at (a) 3 days and (b) 7 days culture. The live cell count exhibits a gradual increase over time. The observable elevation in the optical density of cells, signifying their proliferation rate over time, affirms the biocompatibility of the device. Furthermore, when cells were introduced into a 96-well plate, it became evident that they were uniformly distributed at each outlet during each time interval, demonstrating the device's capability to ensure consistent and even cell distribution. FIG. 7(c) displays MTS assay results for different culture times, and FIG. 7(d) is a photograph of the 8-port device after delivering the cells.

[0031] The incubation period was extended to 14 days while scaling up to the 16-port device. HDFa were cultivated on gelatin hydrogel scaffolds to observe their growth processes after injection through the microfluidic device. Within an hour, cells quickly attached to the gelatin hydrogel surface. On the first day of the 14-day growth period, cellson the patterned scaffold took on an aligned and elongated morphology when their density was sufficient (over 106 / mL) and retained it. The number of live cells was found to have significantly increased over the incubation period, with more uniform cell attachments observed at each port. The MTS proliferation assay demonstrated that cells proliferated successfully over time, with the cell proliferation rate almost doubling from day 7 to day 14. Each outlet’s proliferation rate matched that of a positive control for 14-day cultured cells. The uniformity of cell delivery was also confirmed by comparingthe optical density of cells at each outlet in 96-well plates. These results indicated that the device was able to maintain uniform cell delivery and high proliferation rates, useful for effective and biocompatible cell-based therapies. The gelatin hydrogel scaffolds facilitated cell attachment and growth, mimicking a natural extracellular matrix environment, important for tissue engineering and regenerative medicine applications. The alignment and elongation of cells on the scaffold indicate that the device can support structured tissue formation.

[0032] Similar investigations were carried out with the 32 and 64-port devices with a 14-day culture period. Again, there was a noticeable increase in cell count over the incubation period, and viability was consistently over 90%. Successful cell proliferation was substantiated through MTS proliferation assays, and the proliferation at each outlet port mirrored that of the positive control. Optical density measurements again affirmed an even distribution of cells, with the proliferation rate again nearly doubling from 7 to 14 days. Similar results were also obtained under these conditions with the 128-port device; notably, there was a substantially uniform distribution of cells across the entire 150 cm2surface area. This uniformity is helpful to ensure that all regions of a wound receive an adequate supply of cells, promoting consistent and efficient healing. These multi-port fluidic wound dressing devices may be used to deliver cells over a period of several weeks, providing an adequate supply forthe re-epithelialization phase in wound healing.

[0033] Further, the multi-port fluidic wound dressing device 100 may be employed to fabricate single or multi-layered skin structures to facilitate wound healing or other tissue engineering applications. This presents an alternative to existing techniques for producingskin structures, such as 3D bioprinting, autologous split-thickness skin grafts, or biomaterial scaffolds. Some existing techniques suffer from limitations, such as the availability of donor skin or the need to transfer a fragile, 3D printed construct to the treatment site.

[0034] Fabrication of multi-layered structures was validated by the construction of HDFa strata using a 32-port version of the device 100. HDFa at a concentration of 106cells / mL were deposited onto a collagen 1 % hydrogel substrate layer using the 32-port version of the device 100, and subsequently cultured for 7 days. FIG. 8 illustrates the outcomes of live / dead staining, demonstrating the distribution of cells within a single layer. The left column of FIG. 8 depicts 2D images of the deposited cell layer at 4X and 10X magnification, while the right column depicts a volume of the layer. Cell attachment and penetration into the hydrogel matrix along the z-axis (700 pm) was observed.

[0035] By comparison, FIG. 9 illustrates the results of live / dead staining on a sample of HDFa cells, again at 106cells / mL, encapsulated within collagen (6% v / v) and deposited on a collagen 1 % hydrogel coated on a petri dish using the 32-port version of the device 100, and cultured for 7 days. The lower two images in FIG. 9 again depict the layer at 4X and 10X magnification. Viewing the layer along the z-axis (upper image, 300 pm), it was observed that the cells were substantially aligned in one layer, and with increased cell confluency, indicating that the interaction between the cells and the extracellular matrix significantly influences cell proliferation.

[0036] FIG. 9 illustrates the outcome of live / dead staining of HDFa cells (106cells / mL) encapsulated in collagen hydrogel (6% v / v), deposited onto collagen 1% hydrogel using the 32-port version of the device 100 at intervals of every 2 days, then cultured for 7 days. The z-axis (300-700 pm) shows the alignment of cells in different layers, with proliferation observed throughout the hydrogel matrix.

[0037] To contrast the stratification of delivered cells using the device 100 compared to conventional 3D bioprinting, HDFa cells at a concentration of 106cells / mL were utilized to create a multi-layered structure on a collagen 1 % coated petri dish by three printingmethods: deposition by pipette as a control, deposition using an 8-port version of the device 100, and a 3D bioprinter. First, to visualize the deposited layers, fluorescein dye (yellow-green) was used to label a first layer of collagen hydrogel at a concentration of 0.01 % w / v, and a second layer was labeled with rhodamine B dye (red) at the same concentration. Layers of stained collagen hydrogel (6% v / v) encapsulating HDFa cells at (106cells / mL) were deposited onto a petri dish containing cells at the same concentration. Cells were fixed and stained with DAPI at 0.01 % v / v after 7 days of incubation, then imaged along the z-axis using a confocal microscope at 10X magnification.

[0038] FIG. 10 illustrates the z-axis of the three samples (upper images) and the sample deposition methods (lower images). FIG. 10(a) illustrates a z-axis of 1 .35 mm of the control sample; FIG.10(b) shows the sample produced using the device 100; and FIG. 10(c) shows the sample produced by 3D bioprinting using a RegenHU™ 3DDiscovery printer. While in (c), the layers labelled with fluorescein and rhodamine B appear merged, cells have migrated into these layers along the z-axis at 0.9 mm. The presence of cells between the layers can be observed.

[0039] FIG. 11 illustrates the z-axis of three further samples again produced using the (a) micropipette (control), (b) device 100, and (c) 3D bioprinter. Initially, HDFa cells were printed as the first layer using each method, followed by second and third layers of HDFa cells encapsulated in collagen hydrogel stained with fluorescein and rhodamine B, respectively, at the same concentrations as above. A further layer of HDFa cells at the same concentration was deposited on top of the samples as a fourth layer. Subsequently, the cells were fixed and stained with DAPI at 0.01% v / v, and the layers were imaged along the z-axis using a confocal microscope. FIG. 11 (a) shows the z-axis of the control sample at 1 .66 mm. FIG. 11 (b) shows the z-axis of the sample produced by the device 100 at 1 .66 m. FIG. 11 (c) shows the z-axis of the sample produced by the 3D bioprinter at 1 .1 mm. All three are accompanied by 2D images of the stained cells at different layers. The results here show a sequential arrangement of red and green-tinted fibroblasts, demonstrating the capability to construct 3D, multi-layer cell structures layer by layer without intermixing between the layers. It can be seen that cell viability between the device 100 and 3Dbioprinter appears to be very similar at the given cell densities and after 7 days culture, showing that the device 100 may produce a very similar result to conventional bioprinting with HDFs, with the advantage that deposition of the layers using the device 100 may be performed in situ. By contrast, a structure produced by a 3D bioprinter must be physically transferred to the treatment site after printing.

[0040] Co-culture of HDFs and human epidermal keratinocytes was also investigated using the device 100. The process of wound healing involves the orchestrated recruitment of various cell types to the wound cite in a specified temporal and spatial sequence; reestablishment of the epithelial layer often aligns with the arrival of dermal fibroblasts, and it is highly probably that the interaction between epidermal keratinocytes and dermal fibroblasts plays an important role in restoring tissue integrity. Previously, it has been reported that co-culturing keratinocytes with fibroblasts can trigger paracrine signaling loops between these two cell types, which may take place in vivo to regulate cellular functions. Further, it has been previously reported that fetal keratinocytes can enhance the proliferation and migration of both fetal and adult fibroblasts. Thus, to fabricate skin tissue layers, on day 1 HDFa cells (106cells / mL) were encapsulated into collagen hydrogel (6% v / v) and deposited onto a collagen 1 % coated petri dish using a micropipette as control, and the 8-port version of the device 100. Subsequently, a second layer consisting of human epidermal keratinocytes (HEKs) (LONZA, Switzerland) at a concentration of 106cells / mL encapsulated into collagen hydrogel (6% v / v) was deposited on day 2 and incubated for 3 days. The resulting layers were then examined with live / dead staining to visualize cell viability, again imaged along the z-axis with a confocal microscope at 4X and 10X magnification. FIG. 12(a) shows the z-axis of the control sample fabricated by micropipette (larger image), with 2D images at 4X (right upper image) and 10X (right lower image). FIG. 12(b) correspondingly shows the sample fabricated using the device 100. The formation of distinct layers can be observed. The results were suggestive that the presence of HEKs fosters proliferation of HDFa cells, as evidenced by the observed cell confluency levels. The inadequate healingfor large area burns and other wounds is caused by inadequate cell migration (fibroblasts, keratinocytes, and epithelial cells arising from mesenchymal tissue)during the proliferation phase; thus, the device 100 may facilitate healing of large area wounds.

[0041] The multi-port fluidic wound dressing device 100 of any of the examples above may be provided with one or more drainage holes to enable removal of wound exudate. Excess biofluids can lead to ongoing inflammation and worsen infections, which can impede the healing process, particularly in heavily infected chronic wounds. FIG. 13(a) illustrates an example 64-port device 100 with a set of six drainage holes 40 of approximately 8 mm diameter extending through the device 100 (circled for visibility in FIG. 13(a)), from the first treatment-site contacting surface of the first layer 10 to the opposing surface. In addition to the outlet ports 15 in the first layer 10, both the network 25 of branching channels and the inlet port 35 are visible through the nearly transparent first layer 10. The drainage holes 40 in this example extend through the body of the device 10 between individual channels of the network 25. Any suitable number of holes of suitable size may be formed through the device 100.

[0042] As shown in FIG. 13(c), an absorbent pad 50 may then be placed over the holes 40 on the opposing surface (i.e., against the third layer 30) to capture exudate by capillary action, thus assisting in maintaining a clean wound environment that is conducive to healing. Additionally, the drainage holes 40 may be filled with absorbent material or have a hydrophilic coatingto enhance fluid absorption. In the particular example of FIG. 13, the absorbent pad 50 is provided with one or more complementary projections 55 that fit within the drainage holes 40, as can be seen in FIGS. 13(b) and (c). These projections 55 may contact the wound to enable passive removal of exudate. The absorbent pad 55 also includes an aperture 57 corresponding to the inlet port 35.

[0043] The absorbent pad 50 may be formed of any suitable material, such as an absorbent, hydrophilic polyurethane sponge. The particular example of FIG. 13 is formed of Synthaderm™ polyurethane foam with an absorbent capacity of about 2 mL / cm3and at least 0.3 cm thick, and was found to be capable of absorbing wound exudate at a minimumrate of 3440 g / m2 / day. As such, the absorbent pad 50 may be easily removed and replaced without adhesion to the underlying treatment site.

[0044] It was found that even though increasingly larger devices 100 required more cell solution to cover the treatment area, resulting in more cells delivered, the clustering and aggregation of cells remained uniform at each outlet port of a given device in view of the consistent cell concentration. This demonstrates that the device may be scaled up in size, and even tailored to different wound sizes and types. There may be practical limits to the density of outlet ports 15 in for a given treatment area size and fabrication material, since the area must accommodate a sufficiently complex network 25 of branching channels to accommodate that number of outlet ports 15 while still permitting fluid flow at the desired rate. The above example devices 100 define the network 25 entirely within a 1 mm or lesser depth within the device 100. In alternative embodiments, the network 25 may be created with larger depth, such that parent channels of the network 25 are proximate to the opposing surface bearing the inlet port 35, with daughter channels located progressively closer to the surface bearing the outlet ports 15. Such a device 100 may be more complex to manufacture and may not provide the same flexibility as the example devices 100 above, but may accommodate a denser arrangement of outlet ports 100.

[0045] The example devices 100 above employ a network 25 of branched channels in an Flchannel configuration. This configuration is conducive to a regular grid or array of outlet ports 15, arranged with regular spacing in two dimensions over the surface of the device 100. This enables reasonably consistent delivery of cells or other therapeutic agents through the device 100, as individual branches may deliver fluids from the inlet port 35 to each of the outlet ports 15 at substantially the same volumetric flow rate. A regular grid may comprise different spacing within rows versus columns of outlet ports 15, or may be equidistant (i.e., a square grid). Alternative channel network configurations may be used, although they may not provide a consistent volumetric flow rate at every outlet port 15 without varying the lengths or cross-sectional areas of some branching channels.Alternatively, if different flow rates at different areas of the same treatment site are desired, a plurality of inlet ports 35 and a corresponding plurality of networks 25 of branchedchannels may be provided in a single device 100. This would require multiple pumps, syringes, tubes, or other equipment attached to the device 100 to deliver fluids through each of the inlet ports 35.

[0046] The device 100 may be fabricated using any suitable means and materials besides the fabrication methods and materials described above. A slippery, infused polymer such as iPDMS is advantageous to minimize adhesion to a wound site.

[0047] Thus, there is provided, in one aspect, a fluidic wound dressing device, comprising an inlet port; and a plurality of outlets in a biocompatible surface of the device, the plurality of outlet ports being substantially uniformly spaced within an area of the biocompatible surface; the inlet port in fluid communication with each outlet port of the plurality of outlet ports via a network of branching channels sized to deliver fluid from the inlet port to each outlet port of the plurality of outlet ports at substantially the same volumetric flow rate.

[0048] In another aspect, there is provided a fluidic wound dressing device, comprising a plurality of outlet ports in a biocompatible surface of the device, the plurality of outlet ports distributed within an area of the biocompatible surface such that at least some outlet ports are located at different distances from a center of the area; and an inlet port in an opposing surface of the device, the inlet port being substantially centrally located with respect to the area; the inlet port in fluid communication with each outlet port of the plurality of outlet ports via a network of branching channels, each branching channel being of substantially equal total length and volume, each branching channel terminating at a corresponding outlet port.

[0049] In further aspects, the fluidic wound dressing device comprises an inlet port disposed in an opposing surface of the device, and is substantially centrally located with respect to the area.

[0050] In other aspects, the total branching channel length between the inlet port and each outlet port of the plurality of outlet ports is substantially the same, and / or each outlet port of the plurality of outlet ports is substantially equally spaced from adjacent outlet ports.The outlet ports may be arranged in a regular grid arrangement which may be a square grid arrangement.

[0051] In still other aspects, device comprises a plurality of inlet ports, each inlet port having a corresponding plurality of outlet ports, the corresponding plurality of outlet ports being substantially uniformly spaced within a corresponding area, each inlet port in fluid communication with each outlet port of the corresponding plurality of outlet ports via a corresponding network of branching channels sized to deliver fluid from the inlet port to each outlet port of the corresponding plurality of outlet ports at substantially the same volumetric flow rate. The outlet ports of all pluralities of outlet ports may be substantially uniformly spaced over the corresponding areas when combined, and / or each branching channel is sized to deliver fluid from the inlet port to the corresponding outlet port at substantially the same volumetric flow rate.

[0052] In further aspects, the device comprises a first layer with the plurality of outlet ports extending therethrough, the first layer comprising the biocompatible surface; a second layerwith the inlet port extending therethrough, the second layer comprisingthe opposing surface; and an intermediate layer defining the network of branching channels sealed to each of the first layer and second layer. The intermediate layer may be formed integrally with either the first layer or the second layer.

[0053] In still further aspects, the diameter of the inlet port is up to about 1 mm and a diameter of each outlet port is up to about 3 mm, and / orthe diameter of each outlet port is up to about 2 mm. The dimensions of the ports and the channels may range in size from the micron to millimeter scale, includingfor example 10 microns to 1 cm. The surface area provided by the fluidic wound dressing device may be up to about 200 cm2.

[0054] In further aspects, the network of branching channels comprises an H-channel network of branching channels, each branching channel of the network terminating at a corresponding outlet port.

[0055] Still further, each branching channel comprises a plurality of segments, and relative cross-sectional areas of the plurality of segments may be determined by Murray's law.Further, the channels may have substantially rectangular cross-sections, the ratio of the cross-sectional area of a daughter segment to the cross-sectional area of an immediate parent segment may be between 60 and 65%, and may be approximately 63%.

[0056] In further aspects, the fluid comprises at least one of collagen, hydrogel, cells, an antimicrobial compound, a saline solution, or a growth factor; or a combination of two or more thereof. The device, or at least the biocompatible surface, may be formed of a silicon silicone polymer, which may be a polydimethylsiloxane or oil-infused polydimethylsiloxane.

[0057] In still further aspects, the device includes least one drainage hole or a plurality of drainage holes extending between the biocompatible surface and an opposing surface of the device. The drainage holes may be located between channels of the network of branching channels.

[0058] In yet a further aspect, the device includes an absorbent layer disposed on the opposing surface to retain exudate passing through the at least one drainage hole. The absorbent layer may comprise at least one projection for fitting within the at least one drainage hole. The absorbent layer may comprise absorbent polyurethane pad.

[0059] There is also provided a method for fabricating a multi-layered skin structure, comprising applying the multi-port fluidic dressing device to a surface, which may be a treatment site such as a wound surface; depositing a first layer of collagen hydrogel comprising human dermal fibroblasts (HDFs) and / or human epidermal keratinocytes (HEKs) through the plurality of outlet ports of the device; and depositing at least a second layer of collagen hydrogel comprising HDFs and / or HEKs through the plurality of outlet ports of the device. The collagen hydrogel may comprise HDFs and / or HEKs encapsulated in collagen between about 1 and 6% v / v, and / or at a concentration of at least 104cells / mL, 105cells / mL, or 106cells / mL. In a further aspect, a plurality of second layers are deposited. In still a further aspect, the collagen hydrogel of adjacent layers comprise different amounts of HDFs and / or HEKs. At least one layer of the structure may comprise both HDFsand HEKs, and / or at least some layers may form alternating layers of collagen hydrogel comprising HDFs and collagen hydrogel comprising HEKs.

[0060] It should be understood that this description is not intended to be limiting, and that the examples contemplated herein include all alternatives, modifications, and equivalents as would be appreciated by the person skilled in the art, and are included within the scope of the accompanying claims. Although the features and elements various examples or embodiments may be described as being in particular combinations, the person of ordinary skill in the art will appreciate which features or elements can be used alone, without the other features and elements of the embodiments, or in various combinations with or without other features and elements disclosed herein. Some steps or acts in a process or method may be reordered or omitted, and features and aspects described in respect of one embodiment may be incorporated into other described embodiments.

Claims

Claims1 . A fluidic wound dressing device, comprising: an inlet port; and a plurality of outlets in a biocompatible surface of the device, the plurality of outlet ports being substantially uniformly spaced within an area of the biocompatible surface; the inlet port in fluid communication with each outlet port of the plurality of outlet ports via a network of branching channels sized to deliver fluid from the inlet port to each outlet port of the plurality of outlet ports at substantially the same volumetric flow rate.

2. The fluidic wound dressing device of claim 1 , wherein the inlet port is disposed in an opposing surface of the device.

3. The fluidic wound dressing device of claim 2, wherein the inlet port is substantially centrally located with respect to the area.

4. The fluidic wound dressing device of claim 1 , wherein a total branching channel length between the inlet port and each outlet port of the plurality of outlet ports is substantially the same.

5. The fluidic wound dressing device of claim 4, wherein each outlet port of the plurality of outlet ports is substantially equally spaced from adjacent outlet ports.

6. The fluidic wound dressing device of claim 1 , wherein the device comprises a plurality of inlet ports, each inlet port having a corresponding plurality of outlet ports, the corresponding plurality of outlet ports being substantially uniformly spaced within a corresponding area, each inlet port in fluid communication with each outlet port of the corresponding plurality of outlet ports via a corresponding network of branching channels sized to deliver fluid from the inlet port to each outlet port of the corresponding plurality of outlet ports at substantially the same volumetric flow rate.

7. The fluidic wound dressing device of claim 6, wherein the outlet ports of all pluralities of outlet ports are substantially uniformly spaced over the corresponding areas when combined.

8. A fluidic wound dressing device, comprising: a plurality of outlet ports in a biocompatible surface of the device, the plurality of outlet ports distributed within an area of the biocompatible surface such that at least some outlet ports are located at different distances from a center of the area; and an inlet port in an opposing surface of the device, the inlet port being substantially centrally located with respect to the area; the inlet port in fluid communication with each outlet port of the plurality of outlet ports via a network of branching channels, each branching channel being of substantially equal total length and volume, each branching channel terminating at a corresponding outlet port.

9. The fluidic wound dressing device of claim 8, wherein each branching channel is sized to deliver fluid from the inlet port to the corresponding outlet port at substantially the same volumetric flow rate.

10. The fluidic wound dressing device of any one of claims 1 to 9, wherein the device comprises a first layer with the plurality of outlet ports extending therethrough, the first layer comprising the biocompatible surface; a second layer with the inlet port extending therethrough, the second layer comprising the opposing surface; and an intermediate layer defining the network of branching channels sealed to each of the first layer and second layer.11 . The fluidic wound dressing device of claim 10, wherein the intermediate layer is formed integrally with either the first layer or the second layer.

12. The fluidic wound dressing device of any one of claims 1 to 9, wherein a diameter of the inlet port is up to about 1 mm and a diameter of each outlet port is up to about 3 mm.

13. The fluidic wound dressing device of claim 12, wherein the diameter of each outlet port is up to about 2 mm.

14. The fluidic wound dressing device of any one of claims 1 to 9, wherein the area is up to about 200 cm2.

15. The fluidic wound dressing device of claim 14, wherein the area is up to about 150 cm2.

16. The fluidic wound dressing device of any one of claims 1 to 9, wherein the plurality of outlet ports is arranged in a substantially rectangular grid array.

17. The fluidic wound dressing device of any one of claims 1 to 9, wherein the network of branching channels comprises an H-channel network of branching channels, each branching channel of the network terminating at a corresponding outlet port.

18. The fluidic wound dressing device of any one of claims 1 to 9, wherein each branching channel comprises a plurality of segments, and wherein relative cross-sectional areas of the plurality of segments are determined by Murray's law.

19. The fluidic wound dressing device of any one of claims 1 to 9, wherein each branching channel comprises a plurality of segments, each segment of the plurality of segments having a substantially rectangular cross-section, and within a branching channel, a ratio of a cross-sectional area of a daughter segment to a cross-sectional area of an immediate parent segment is between 60 and 65%.

20. The fluidic wound dressing device of claim 19, wherein the ratio is approximately 63%.21 . The fluidic wound dressing device of any one of claims 1 to 7 or 9, wherein the fluid comprises a hydrogel.

22. The fluidic wound dressing device of any one of claims 1 to 7 or 9, wherein the fluid comprises at least one of collagen, hydrogel, cells, an antimicrobial compound, a saline solution, or a growth factor; or a combination of two or more thereof.

23. The fluidic wound dressing device of any one of claims 1 to 9, wherein at least the biocompatible surface of the device is formed of a silicone polymer.

24. The fluidic wound dressing device of claim 23, wherein the silicone polymer is a polydimethylsiloxane.

25. The fluidic wound dressing device of claim 24, wherein the polydimethylsiloxane is an oil-infused polydimethylsiloxane.

26. The fluidic wound dressing device of any one of claims 1 to 9, further comprising at least one drainage hole extending between the biocompatible surface and an opposing surface of the device.

27. The fluidic wound dressing device of claim 26, wherein the at least one drainage hole is located between channels of the network of branching channels.

28. The fluidic wound dressing device of claim 26, further comprising a plurality of drainage holes extending between the biocompatible surface and an opposing surface of the device.

29. The fluidic wound dressing device of claim 26, further comprising an absorbent layer disposed on the opposing surface to retain exudate passing through the at least one drainage hole.

30. The fluidic wound dressing device of claim 29, wherein the absorbent layer comprises at least one projection forfitting within the at least one drainage hole.31 . The fluidic wound dressing device of claim 29, wherein the absorbent layer comprises an absorbent polyurethane pad.

32. A method for fabricating a multi-layered skin structure, comprising: applying the device of any one of claims 1 to 9 to a surface;depositing a first layer of collagen hydrogel comprising human dermal fibroblasts (HDFs) and / or human epidermal keratinocytes (HEKs) through the plurality of outlet ports of the device; and depositing at least a second layer of collagen hydrogel comprising HDFs and / or HEKs through the plurality of outlet ports of the device.

33. The method of claim 32, wherein the surface is a wound surface.

34. The method of claim 32, wherein the collagen hydrogel comprises HDFs and / or HEKs encapsulated in collagen between about 1 and 6% v / v.

35. The method of claim 32, wherein the collagen hydrogel comprises HDFs and / or HEKs encapsulated in collagen at a concentration of at least 104cells / mL, 105cells / mL, or 106cells / mL.

36. The method of claim 32, wherein depositing the at least the second layer of the collagen hydrogel comprises depositing a plurality of second layers of the collagen hydrogel.

37. The method of claim 32, wherein the collagen hydrogel of adjacent layers comprise different amounts of HDFs and / or HEKs.

38. The method of claim 32, wherein at least one of the first layer or at least the second layer of collagen hydrogel comprises both HDFs and HEKs.

39. The method of claim 32, wherein the first layer and the at least the second layer of collagen form alternating layers of collagen hydrogel comprising HDFs and collagen hydrogel comprising HEKs.

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