While three-dimensional (3-D) cell culture systems bet- ter mimic the microstructure of intact organs relative to 2-D cultures, the 3-D systems still fail to fully recapitulate organ-level physiologic functions presumably due to an in- ability to fully control the microenvironment of the organoid. Consequently, a growing trend is to build ‘organ-on-chip’ devices which integrate the 3-D tissue cul- ture systems with microdevice technologies to offer enhanced control of both surface and fluidic conditions [13-16]. However due to the difficulty in obtaining and isolating primary tissue, these ‘organ-on-chip’ devices often utilize tumor cell lines which are incapable of demon- strating organ-level physiologic function. For example, ‘gut- on-chips’ devices are frequently assembled by placing Caco-2 tumor cells within microdevices [17-19]. The Caco- 2 tumor cell line has been adapted for tissue culture and poorly mimics the intestinal epithelium in terms of architecture, growth factor response, differentiation, gene expression and susceptibility to apoptosis [20,21]. A sig- nificant challenge to the ‘organ-on-chip’ community is the development of optimized strategies to isolate high-quality primary cells for culture within a microdevice.
Very dynamic R&D activities have been translated into technological advances in both microfluidics and tissue engineering, providing OoCs with key added values towards a more accurate view of what hap- pens in humans compared to other existing models (Fig. 3). OoC technology allows reconstitution of the microarchitecture of the organ supported by the design of a dedicated mechanical context matching the shape, surface pattern and stiffness of organ-specific microenvironments. Precise microfluidic flow control enables optimal oxygenation and nutrition that not only afford long-term viability of healthy tissues, but also an efficient circulation of immune system cells, antibodies, biochemical signalling molecules and me- tabolites, and the ability to collect tiny secretion volumes for analyses. Continuous perfusion and mechan- ical stress help to build dynamic tissue models, supposedly far more relevant than conventional static cell cultures, and enable the control of spatiotemporal chemical gradients and mechanical cues to study the influence of the microenvironment on the cells. Moreover, OoCs allow the precise investigation of specific tissue-tissue interfaces and biological events that cannot be monitored in animals or human patients. Their minute size, the control of the microenvironment, the facilitated optical access at high spatial resolution, and the integration of biosensors for real-time data collection bring OoCs large advantages over other mod- els 1 . The tightly monitored regulation of the cellular environment and homeostasis should facilitate long-
Although it has been well established that exposure to LPS induces cytokine responses [16, 32], and in the case of other organ systems that LPS exposure has been linked to re- duced tight junction protein expression , relatively little is known about its effects on the BBB and how this com- pares to the direct cytokine exposure that LPS is supposed to induce. To study the effects of inflammatory signals on BBB function, we leveraged novel microfluidic technology in a dual-chambered device, creating a system that contains the relevant cell types for BBB formation and enables these cells to form the barrier in the presence of flow and a differ- ential serum concentration from the vascular to the brain side of the device . Once the brain-derived endothelial cells, astrocytes, and pericytes have had a chance to form the basis for the BBB, hiPSC-derived cortical neurons and co-differentiating astrocytes are suspended in a collagen gel and loaded on top of the astrocytes and pericytes. This en- tire component is what we call the NeuroVascular Unit (NVU), which previously has been shown to restrict diffu- sion of small molecules and facilitate active transport . Having established our cellular model of the BBB (Fig. 1a), we then sought a concentration of LPS and cytokine cock- tail that, while activating the system, would not cause cell death. We wanted the highest concentration that did not increase cell death from control. Over the dose ranges tested (25 to 100 μg/ml), we found that 100 μg/mL showed no cell death in either the vascular chamber containing endothelial cells (Fig. 1b) or in the brain chamber contain- ing pericytes, astrocytes, and neurons (Fig. 1c), as assessed by live/dead staining after 24 h of LPS or cocktail exposure.
Microfluidic organ-on-chip approaches offer a precise means to control tissue composition and architecture in an in vitro 3D microdevice that further incorporates vascular perfusion and micro- biofabrication (Takebe et al., 2017). This approach gives researchers finer control over multiple physiological phenomena such as tissue- tissue interactions and physicochemical niche cues, as well as physical forces that occur in living organs, such as breathing movement, shear stress, peristalsis and tension (Bhatia and Ingber, 2014). In 2010, a human lung-on-chip was created using classical soft lithography and microfluidic devices to reconstitute the functional alveolar-capillary interface of the lung (Huh et al., 2010). Since this discovery, organ-on-chip systems have been applied in iPSC disease modeling for an increasingly wide range of different diseases, including Barth syndrome-associated cardiomyopathy, drug-induced kidney glomerular injury, blood- brain barrier function and skin wound healing (Low and Tagle, 2017).
Abstract- Continuous scaling of CMOS technology makes it possible to integrate a large number of heterogeneous devices that need to communicate efficiently on a single chip. For this efficient routers are needed to takes place communication between these devices. This paper gives the design of on-chip routers based on optimizing power consumption and chip area. Proposed architecture of on-chip router in this paper give the results in which power consumption is reduced and silicon area is also minimize.
Figure 3 shows the optimization plot for chip thickness. From the plot below, the optimal cutting conditions for chip thickness was found to be at lower cutting speed (765 rpm), lower feed (31.5 mm/rev) and lower axial depth of cut (0.50 mm). The optimal values are displayed in red on the plot. At these machining conditions, it is expected that one would produce a chip with small thickness. The objective of the optimization process is to minimize the chip thickness. When the chip is thin, the saw tooth becomes smaller which thus, reduces machining instability thereby improving the tool life. The composite desirability is 1.
Deepa N P, obtained her Bachelor’s degree in Electronics and Communication Engineering from Kuvempu University, Karnataka, India and Masters in Digital Communication Engineering from Visvesvaraya Technological University, Belagavi, Karnataka, India. She is a part time research scholar in the field of short range wireless communication and has 15 years of teaching experience in various engineering colleges. She has guided many post graduate and undergraduate students for their final year projects. She has published research papers in various journals and national / international conferences. Her area of interest includes Wireless communication, Network on Chip, RF and Microwave, Signal Processing and Image processing.
diameter of swollen gel at each temperature t respectively. The environmental temperature was controlled by a thermo plate (TOKAI HIT. Co. Ltd., TP-CH110R) which is placed under the microfluidic chip. The photo of the shape of gel pattern was taken by the vision sensor attached to the microscope. We waited 10 minutes to confirm stabilization of each temperature for experiment. We measured the shape of patterned gel during temperature increasing phase, and then decreasing phase. The expansion rate is defined as the ratio between the swollen diameter at each temperature condition and the patterned diameter of the fabricated pillar. Figure 4 shows how Bioresist patterns change. The pattern diameter and the height are 20 μm and 4 μm, respectively. From Figures 4(b) through (c), we can see that the shape of Bioresist is swollen at the low temperature condition, and that the shape shrinks with high temperature. Figure 5 shows the relationship between the expansion rate and
Capillary electrophoresis (CE) suffers from a relatively small sensitivity—at least in case of optical detection transversely to the capillary axis due to the small capillary inner diameters in the range of 50 - 100 µm. Different concepts like bubble, U-, or Z-cells have been used to tackle that problem already in the nineties of the last century. But the U- and Z-cells have typically been extra cells with larger inner channel diameters and no optimization for optical waveguiding and the bubble cell per se did not allow for optical waveguiding. In the case of on-chip capillary electrophoresis (chip-CE) a U-cell can be implemented quite easily on the chip. Here we show how leaky optical waveguiding can be employed to improve optical detection. Proper U-channel design and prepa- ration by wet-chemical etching of the fused silica sub- and superstrate, making the U-channel bend a part of the optical input lens system, can help to achieve high coupling efficiency with loss coeffi- cients around 2 dB and low waveguiding loss.
In 2013, Woodruff’s team created the first “organ-on-a-chip” model that functionally re-creates in vitro the entire 28-days menstrual cycle, where human and mouse cells from several reproductive organs were grown in a network of tiny and interconnected microengineered units . Three different microfluidic platforms (MFP) termed Solo-MFP, Duet-MFP and Quintet-MFP were developed and used as a platform able to sustain tissue-level functions for the length of the human menstrual cycle. The Solo-MFP and Duet-MFP systems were designed for single tissue culture and are based on pneumatic actuation technology that uses sequential application of pressure and vacuum to valve and pump membranes to enable fluid to move throughout the system. The Quintet-MFP was designed for multiple tissue culture and is based on embedded electromagnetic actuation technology which uses a series of micropumps and individually controllable actuators that enable precise flow control. Thus, the use of tubes, valves and pumps allows to push air and fluids through the system, mimicking the body’s natural circulation.
In recent years, the flip chip packaging technology has been developed rapidly. There are many scholars have carried out lots of research on flip chip packaging technology domestic and overseas. The research is mainly concentrated on the packaging process, material and reliability, etc. In terms of the flip chip bonding mechanism, there is much research has been performed about the thermal ultrasonic flip chip bonding of gold bump, especially the research made by professor Lei Hanand Junhui Li in Central South University. Their team has done a lot of work on ultrasonic bonding mechanism, micro structure generation in bonding region, multi interface transfer and transformation rule, process design and monitoring system. They have made a significant contribution to the study of ultrasonic bonding theory. However, there is little research on the flip chip bonding process of reflow soldering at present. Therefore, this paper studies the flip chip bonding of reflow soldering by simulation and experiment, and it can provide a theoretical basis for the development of high density flip chip bonding equipment.
With the emergence of Internet of Things and information revolution, the demand of high performance computing systems is increasing. The copper interconnects inside the computing chips have evolved into a sophisticated network of interconnects known as Network on Chip (NoC) comprising of routers, switches, repeaters, just like computer networks. When network on chip is implemented on a large scale like in Multicore Multichip (MCMC) systems for High Performance Computing (HPC) systems, length of interconnects increases and so are the problems like power dissipation, interconnect delays, clock synchronization and electrical noise. In this thesis, wireless interconnects are chosen as the substitute for wired copper interconnects. Wireless interconnects offer easy integration with CMOS fabrication and chip packaging. Using wireless interconnects working at unlicensed mm-wave band (57-64GHz), high data rate of Gbps can be achieved.
The first problem is to connect the chip under test to the probe chip and keep their faces as close together as possible, ideally touching. This was depicted previously in Figure 3.2. The chips connect through the usual solder bump interconnections the chip under test will ulti- mately employ for bonding to the MCM substrate. However, to achieve the second require- ment pits must be etched into the probe chip to accept the bumps on the test chip and keep the faces close. If the normal size solder bumps of around 100 µm diameter are used then the craters would have to be huge, over 50 µm deep. Such large connections are not really neces- sary though, since they do not have to pass the normal reliability criteria as they will only be used for chip debug and not in normal service. Therefore, smaller bumps, call them micro bumps, of around 20 µm diameter can be used. This only requires a depth of something over 10 µm. These holes have insulator grown, a normal signal conductor metal and a wettable metals deposited to form the receptors for the solder bumps. The chip under test only needs smaller overglass cuts to reduce the bump size. The mask changes for the chip under test are, therefore, small.
Recall the design of 6-point stencil using the loop or array order has six input FIFOs that receive six elements for computing an output element. Each FIFO receives the appropriate data with the FIFO select control signal. The control signal specifies which of the six FIFOs the requested element from the memory interface will be stored into. When the address generator requests the elements numerically, it is not feasible to have a large decoder to decode the location of each element to know which FIFO the element will be stored into. So all the requested elements have to be stored in the same on-chip buffer. On the other hand, if all the elements are stored in one buffer with a single read port, then it would take six cycles to read six elements for computing an output element, decreasing the computation throughput of the kernel and thus requiring a very large buffer to store the elements.
area, especially tumor tissue, can be confined and optimized. The trapped living cells/drug can be transported into the liquid core channels/capillaries via specifically designed optical tweezers and waveguides, in which drug molecules/ cells can be trapped, stored, and delivered to specific targets. Such a system can be fabricated on a chip and used as a drug injection tool (syringe) for targeted treatment to cancerous tissue. However, the size of the drug molecule is shown to be varied by the character of the interacting light used, ie, a standing wave versus a propagating wave. 18 The design
Designers of medical implants face three primary challenges: size, cost and power consumption. At the same time, there is a desire for an increase in the capability of these implants – both in terms of an expansion in the scale of current functionality, such as increasing the number of electrodes in neural recording or retinal prosthesis implants, as well as adding new functionality. Size and power considerations have driven the use of specialized, highly-integrated system-on-chip designs (e.g. , , , ), which can be cost-prohibitive for the low-volume applications typical in the biomedical market. Additionally, as described in Section 1.3, increasing the scale of existing functionality can make even highly-integrated designs, such as a thousand-electrode retinal prosthesis, approach the limits of implant size in small and delicate organs such as the eye. Given the low leakage power and high voltage tolerances typically required of biomedical implants, it seems unlikely that CMOS scaling alone can provide the size reductions required to meet future demands for increased capability; even if it can, increased development complexity and mask costs will certainly exacerbate already steep development costs and long times-to- market.
Conversion of 2D NoC into three dimensional 3D NoC has been proposed and the 3D Networks-on-Chip (3D NoCs) have been attracted an interest to solve on-chip communication demands for future multipurpose systems. In this paper, a brief idea of 3D NoCs optimization techniques of modeling and evaluation of alternate Noc topologies, routing algorithms and mapping techniques are presented to achieve optimized area and power.
In 1991, a national workshop on increasing organ donation convened by Surgeon General Antonia Novella concluded that “stronger efforts were needed to ensure compliance with existing routine inquiry and required request laws….. Unfortunately, most State laws contain neither provisions nor monies to assure adequate compliance” . In 1998, HCFA published a final rule for organ, tissue and eye donation (63 Federal Register 33875) as part of the revised Medicare and Medicaid Conditions of Participation for Hospitals (42 CFR Part 482). To facilitate best practices for increasing organ donation, two key requirements were added to the conditions of participation. First, hospitals must refer imminent death and deceased patients to their OPO in a timely manner. Second, only OPO staff or a trained hospital staff may approach families about organ donation. For the critical care staff, the new requirements were a departure from their traditional role. Lacking experience on this front, many hospitals struggled to implement the referral and request requirements.
The most distinctive result generated by this study is the high preference of SVA repeats for BORIS binding, as com- pared to binding by CTCF in K562 (Fig. 4). Unfortunately, in the absence of ChIP data for BORIS from human testis one cannot be absolutely sure that it is also the situation in normal testis. The functions for SVA that are described so far are attributed to the disruption/features of insertion sites rather than to the transcription originating within the insertion [103, 104]; yet the finding of BORIS binding hints at the regulatory role of SVA VNTRs themselves. The presence of several BORIS binding sites within the VNTR repeats (Figs. 4c, f, 5), which are actually required for SVA transposition , indicates that the BORIS protein and SVA elements may have even undergone co-evolution, as has been recently suggested for other ZF proteins . Thus, one may expect the SVA elements to play a nota- ble regulatory role in germline development and genome evolution in primates. In that regard, the recent studies on gibbon genome [2, 105] provided some invaluable insight into the new level of plasticity that SVA-like elements LAVA infused into primate genomes. At present, one can- not conclude whether SVA TEs merely represent a genetic load or actually have a physiological role in germline. Despite human SVAs being associated with at least some chromosomal breaks , we could probably exclude the direct contribution of SVA elements into the meiotic recombination, as DSB maps of human meiosis  did not correspond to SVA locations (not shown).