Print Your Skin And Organs Using 3D Bioprinter

3D Bioprinting , An On-demand Printing Of Natural Skin And Organs

3D bioprinting is the process of generating spatially-controlled cell patterns using 3D printing technologies, where cell function and viability are preserved within the printed construct.^[1]:1 The first patent related to this technology was filed in the United States in 2003 and granted in 2006

Using 3D bioprinting for fabricating biological constructs typically involves dispensing cells onto a biocompatible scaffold using a successive layer-by-layer approach to generate tissue-like three-dimensional structures. Given that every tissue in the body is naturally compartmentalized of different cell types, many technologies for printing these cells vary in their ability to ensure stability and viability of the cells during the manufacturing process. Some of the methods that are used for 3D bioprinting of cells are photolithography, magnetic bioprinting, stereolithography, and direct cell extrusion. When a bioprinted pre-tissue is transferred to an incubator then this cell-based pre-tissue matures into a tissue.

While most are familiar with the potential for 3D printers to pump out plastic odds and ends for around the home, the technology also has far-reaching applications in the medical field. Research is already underway to develop 3D bioprinters able to create things as complex as human organs, and now engineering students in Canada have created a 3D printer that produces skin grafts for burn victims.

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The new machine was developed by University of Toronto engineering students Arianna McAllister and Lian Leng, who worked in collaboration with Professor Axel Guenther, Boyang Zhang and Dr. Marc Jeschke, the head of Sunnybrook Hospital's Ross Tilley Burn Centre.

While the traditional treatment for serious burns involves removing healthy skin from another part of the body so it can be grafted onto the affected area, the PrintAlive machine could put an end to such painful harvesting by printing large, continuous layers of tissue – including hair follicles, sweat glands and other human skin complexities – onto a hydrogel. Importantly, the device uses the patient's own cells, thereby eliminating the problem of the tissue being rejected by their immune system.

Because growing a culture of a patient's skin cells ready for grafting can typically take more than two weeks, the machine prints the patient's cells out in patterns of spots or stripes rather than a continuous sheet, to make them go further. The result is a cell-populated wound dressing that reproduces key features of human skin and can be precisely controlled in terms of thickness, structure and composition.

A typical process for bioprinting 3D tissues

Having been under development since 2008, the team recently completed a second-generation, pre-commercial prototype that they say is smaller than an average microwave. This makes it portable enough to easily transport, which gives it the potential to one day revolutionize burn care in rural and developing areas around the world.

 Examples of human-scale bioprinted tissues

"Ninety per cent of burns occur in low and middle income countries, with greater mortality and morbidity due to poorly-equipped health care systems and inadequate access to burn care facilities," says Jeschke. "Regenerating skin using a patient’s own stem cells can significantly decrease the risk of death in developing countries."

Time Line For Developing Different kinds of Bio printers

Inkjet bioprinting. 

Inkjet printers (also known as drop-on-demand printers) are the most commonly used type of printer for both nonbiological and biological applications. Controlled volumes of liquid are delivered to predefined locations. The first inkjet printers used for bioprinting applications were modified versions of commercially available 2D ink-based printers. The ink in the cartridge was replaced with a biological material, and the paper was replaced with an electronically controlled elevator stage to provide control of the z axis. (the third dimension in addition to the x and yaxes). Now, inkjet-based bioprinters are custom-designed to handle and print biological materials at increasing resolution, precision and speed. Inkjet printers use thermal. or acoustic forces to eject drops of liquid onto a substrate, which can support or form part of the final construct.

Thermal inkjet printers function by electrically heating the print head to produce pulses of pressure that force droplets from the nozzle. Several studies have demonstrated that this localized heating, which can range from 200 °C to 300 °C, does not have a substantial impact either on the stability of biological molecules, such as DNA., or on the viability or post-printing function of mammalian cells. It has been demonstrated that the short duration of the heating (~2 μs) results in an overall temperature rise of only 4–10 °C in the printer head⁵⁵. The advantages of thermal inkjet printers include high print speed, low cost and wide availability. However, the risk of exposing cells and materials to thermal and mechanical stress, low droplet directionality, nonuniform droplet size, frequent clogging of the nozzle and unreliable cell encapsulation pose considerable disadvantages for the use of these printers in 3D bioprinting.

Many inkjet printers contain a piezoelectric crystal that creates an acoustic wave inside the print head to break the liquid into droplets at regular intervals. Applying a voltage to a piezoelectric material induces a rapid change in shape, which in turn generates the pressure needed to eject droplets from the nozzle. Other inkjet printers use an acoustic radiation force associated with an ultrasound field to eject liquid droplets from an air-liquid interface. Ultrasound parameters, such as pulse, duration and amplitude, can be adjusted to control the size of droplets and the rate of ejection. Advantages of acoustic inkjet printers include the capability to generate and control a uniform droplet size and ejection directionality as well as to avoid exposure of cells to heat and pressure stressors. Additionally, the sheer stress imposed on cells at the nozzle tip wall can be avoided by using an open-pool nozzle-less ejection system. This reduces the potential loss of cell viability and function, and avoids the problem of nozzle clogging. Acoustic ejectors can be combined as multiple ejectors in an adjustable array format, facilitating simultaneous printing of multiple cell and material types. Even so, there remain some concerns regarding the 15–25 kHz frequencies used by piezoelectric inkjet bioprinters and their potential to induce damage of the cell membrane and lysis. Inkjet bioprinters also have limitations on material viscosity (ideally below 10 centipoise) owing to the excessive force required to eject drops using solutions at higher viscosities.

One common drawback of inkjet bioprinting is that the biological material has to be in a liquid form to enable droplet formation; as a result, the printed liquid must then form a solid 3D structure with structural organization and functionality. Our group and others have shown that this limitation could be addressed by using materials that can be crosslinked after deposition by printing using chemical, pH or ultraviolet mechanisms. However, the requirement for crosslinking often slows the bioprinting process and involves chemical modification of naturally occurring ECM materials, which changes both their chemical and material properties. Additionally, some crosslinking mechanisms require products or conditions that are toxic to cells, which results in decreased viability and functionality. Another limitation encountered by users of inkjet-based bioprinting technology is the difficulty in achieving biologically relevant cell densities. Often, low cell concentrations (fewer than 10 million cells/ml). are used to facilitate droplet formation, avoid nozzle clogging and reduce shear stress. Higher cell concentrations may also inhibit some of the hydrogel crosslinking mechanisms.

Notwithstanding these drawbacks, inkjet-based bioprinters also offer advantages, including low cost, high resolution, high speed and compatibility with many biological materials. Another advantage of inkjet printing is the potential to introduce concentration gradients of cells, materials or growth factors throughout the 3D structure by altering drop densities or sizes. Because of the availability of standard 2D inkjet printers, researchers in many labs can readily access, modify and experiment with 3D inkjet–based bioprinting technology. Commercially available inkjet bioprinters are also relatively cost-effective owing to their simple components and readily available design and control software. The wide application of this technology by many groups has accelerated advances in the capacity of inkjet bioprinters to accurately deposit with high resolution and precision controllable droplet sizes with uniform cellular densities. Droplet size and deposition rate can be controlled electronically, and can range from <1 pl to >300 pl in volume with rates of 1–10,000 droplets per second. Patterns of single drops, each containing one or two cells, in lines ~50 μm wide, have been printed. Future advances will continue to adapt this technology to handle and deposit other biologically relevant materials, in a manner that both facilitates their printing and provides the essential biological, structural and functional components of the tissue. Additional complexities, such as the requirement for multiple cell types and materials, will also have to be addressed.

Notable examples of the inkjet bioprinting approach include the regeneration of functional skin and cartilage in situ. The high printing speed of the approach enables direct deposition of cells and materials directly into skin or cartilage lesions. These applications achieved rapid crosslinking of the cell-containing material via either a biocompatible chemical reaction or a photoinitiator and crosslinking through exposure of the material to ultraviolet light. The inkjet approach facilitated the deposition of either primary cells or stem cell types with uniform density throughout the volume of the lesion, and maintained high cell viability and function after printing. These studies demonstrate the potential of inkjet-based bioprinting to regenerate functional structures.

Layered cartilage constructs have also been fabricated in vitro using a combination of electrospinning and inkjet bioprinting. The hybrid electrospinning–inkjet bioprinting technique enabled the fabrication of a layered construct that supported cell function and maintained suitable mechanical and structural properties. Inkjet bioprinters have also been used to fabricate bone constructs, matured in vitro before implantation into mice. These constructs continued to maturein vivo and formed highly mineralized tissues with similar density as endogenous bone tissue.

Microextrusion bioprinting. 

The most common and affordable nonbiological 3D printers use microextrusion. Microextrusion bioprinters usually consist of a temperature-controlled material-handling and dispensing system and stage, with one or both capable of movement along the x, yand z axes, a fiberoptic light source to illuminate the deposition area and/or for photoinitiator activation, a video camera for x-y-z command and control, and a piezoelectric humidifier. A few systems use multiple print heads to facilitate the serial dispensing of several materials without retooling. Nearly 30,000 3D printers are sold worldwide every year, and academic institutions are increasingly purchasing and applying microextrusion technology in tissue and organ engineering research. Industrial printers are considerably more expensive but have better resolution, speed, spatial controllability and more flexibility in the material they can print.

Microextrusion printers function by robotically controlled extrusion of a material, which is deposited onto a substrate by a microextrusion head. Microextrusion yields continuous beads of material rather than liquid droplets. Small beads of material are deposited in two dimensions, as directed by the CAD-CAM software, the stage or microextrusion head is moved along the z axis, and the deposited layer serves as a foundation for the next layer. A myriad of materials are compatible with microextrusion printers, including materials such as hydrogels, biocompatible copolymers and cell spheroids. The most common methods to extrude biological materials for 3D bioprinting applications are pneumatic or mechanical (piston or screw) dispensing systems. Mechanical dispensing systems might provide more direct control over the material flow because of the delay of the compressed gas volume in pneumatic systems. Screw-based systems might give more spatial control and are thought to be beneficial for the dispensing of hydrogels with higher viscosities, although pneumatic systems could also be suited to dispense high-viscosity materials. Pneumatically driven printers have the advantage of having simpler drive-mechanism components, with the force limited only by the air-pressure capabilities of the system. Mechanically driven mechanisms have smaller and more complex components, which provide greater spatial control but often at reduced maximum force capabilities.

Microextrusion methods have a very wide range of fluid properties that are compatible with the process, with a broad array of biocompatible materials described in the literature. Materials with viscosities ranging from 30 mPa/s to >6 × 10⁷ mPa/s  have been shown to be compatible with microextrusion bioprinters, with higher-viscosity materials often providing structural support for the printed construct and lower-viscosity materials providing a suitable environment for maintaining cell viability and function. For microextrusion bioprinting, researchers often exploit materials that can be thermally crosslinked and/or possess sheer-thinning properties. Several biocompatible materials can flow at room temperature, which allows their extrusion together with other biological components, but crosslink into a stable material at body temperature. Alternatively, materials that flow at physiologically suitable temperatures (35–40 °C), but crosslink at room temperature may also be useful for bioprinting applications. Materials with shear-thinning properties are commonly used for microextrusion applications. This non-newtonian material behavior causes a decrease in viscosity in response to increases in shear rate. The high shear rates that are present at the nozzle during biofabrication allow these materials to flow through the nozzle, and upon deposition, the shear rate decreases, causing a sharp increase in viscosity. The high resolution of microextrusion systems permits the bioprinter to accurately fabricate complex structures designed using CAD software and facilitate the patterning of multiple cell types.

The main advantage of microextrusion bioprinting technology is the ability to deposit very high cell densities. Achieving physiological cell densities in tissue-engineered organs is a major goal for the bioprinting field. Some groups have used solutions comprised only of cells to create 3D tissue constructs with microextrusion printing. Multicellular cell spheroids are deposited and allowed to self-assemble into the desired 3D structure. Tissue spheroids are thought to possess material properties that can replicate the mechanical and functional properties of the tissue ECM. Depending on the viscoelastic properties of the building blocks, the apposed cell aggregates fuse with each other, forming a cohesive macroscopic construct. One advantage of the self-assembling spheroid strategy is potentially accelerated tissue organization and the ability to direct the formation of complex structures. This approach shows promise in enabling the generation of an intraorgan branched vascular tree in 3D thick tissue or organ constructs by patterning self-assembling vascular tissue spheroids, in 3D bioprinted organs. The most common technology used for scaffold-less tissue spheroid bioprinting is mechanical microextrusion.

Cell viability after microextrusion bioprinting is lower than that with inkjet-based bioprinting; cell survival rates are in the range of 40–86%, with the rate decreasing with increasing extrusion pressure and increasing nozzle gauge. The decreased viability of cells deposited by microextrusion is likely to result from the shear stresses inflicted on cells in viscous fluids. Dispensing pressure may have a more substantial effect on cell viability than the nozzle diameter. Although cell viability can be maintained using low pressures and large nozzle sizes, the drawback may be a major loss of resolution and print speed. Maintaining high viability is essential for achieving tissue functionality. Although many studies report maintenance of cell viability after printing, it is important for researchers to demonstrate that these cells not only survive, but also perform their essential functions in the tissue construct.

Increasing print resolution and speed is a challenge for many users of microextrusion bioprinting technology. Nonbiological microextrusion printers are capable of 5 μm and 200 μm resolution at linear speeds of 10–50 μm/s. Whether these parameters can be matched using biologically relevant materials while maintaining high cell viability and function is yet to be seen. Use of improved biocompatible materials, such as dynamically crosslinked hydrogels, that are mechanically robust during printing and that develop secondary mechanical properties after printing might help to maintain cell viablity and function after printing. Single-phase, dual-phase and continuous-gradation scaffolds are also being designed using similar principles. Additionally, improvements in nozzle, syringe or motor-control systems might reduce print times as well as allow deposition of multiple diverse materials simultaneously.

Microextrusion bioprinters have been used to fabricate multiple tissue types, including aortic valves, branched vascular trees and in vitro pharmokinetic as well as tumor models. Although the fabrication time can be slow for high-resolution complex structures, constructs have been fabricated that range from clinically relevant tissue sizes down to micro-tissues in microfluidic chambers.

Laser-assisted bioprinting. 

Laser-assisted bioprinting (LAB) is based on the principles of laser-induced forward transfer. Initially developed to transfer metals, laser-induced forward transfer technology has been successfully applied to biological material, such as peptides, DNA and cells. Although less common than inkjet or microextrusion bioprinting, LAB is increasingly being used for tissue- and organ-engineering applications. A typical LAB device consists of a pulsed laser beam, a focusing system, a 'ribbon' that has a donor transport support usually made from glass that is covered with a laser-energy-absorbing layer (e.g., gold or titanium) and a layer of biological material (e.g., cells and/or hydrogel) prepared in a liquid solution, and a receiving substrate facing the ribbon. LAB functions using focused laser pulses on the absorbing layer of the ribbon to generate a high-pressure bubble that propels cell-containing materials toward the collector substrate.

The resolution of LAB is influenced by many factors, including the laser fluence (energy delivered per unit area), the surface tension, the wettability of the substrate, the air gap between the ribbon and the substrate, and the thickness and viscosity of the biological layer. Because LAB is nozzle-free, the problem of clogging with cells or materials that plague other bioprinting technologies is avoided. LAB is compatible with a range of viscosities (1–300 mPa/s) and can print mammalian cells with negligible effect on cell viability and function. LAB can deposit cells at a density of up to 10⁸ cells/ml with microscale resolution of a single cell per drop using a laser pulse repetition rate of 5 kHz, with speeds up to 1,600 mm/s .

Despite these advantages, the high resolution of LAB requires rapid gelation kinetics to achieve high shape fidelity, which results in a relatively low overall flow rate. Preparation of each individual ribbon, which is often required for each printed cell or hydrogel type, is time-consuming and may become onerous if multiple cell types and/or materials have to be co-deposited. Because of the nature of the ribbon cell coating, it can be difficult to accurately target and position cells. Some of these challenges might be overcome by using cell-recognition scanning technology to enable the laser beam to select a single cell per pulse. This so-called 'aim-and-shoot' procedure could ensure that each printed droplet contains a predefined number of cells. However, statistical cell printing can be achieved using a ribbon with very high cell concentrations, avoiding the need for such specific cell targeting. Finally, metallic residues are present in the final bioprinted construct, owing to the vaporization of the metallic laser-absorbing layer during printing. Approaches to avoid this contamination include the use of nonmetallic absorbing layers and modifying the printing process to not require an absorbable layer. The high cost of these systems is also a concern for basic tissue-engineering research, although as is the case with most 3D printing technologies, these costs are rapidly decreasing.

The application of LAB to fabricate a cellularized skin construct demonstrated the potential to print clinically relevant cell densities in a layered tissue construct, but it is unclear whether this system can be scaled up for larger tissue sizes. In vivo LAB has been used to deposit nano-hydroxyapatite in a mouse calvaria 3D defect model. In these studies, a 3 mm diameter, 600 μm–deep calvarial hole was filled as a proof of concept. Laser 3D printing has been used to fabricate medical devices, such as a customized, noncellular, bioresorbable tracheal splint that was implanted into a young patient with localized tracheobronchomalacia. Future studies might use materials that can directly integrate into a patient's tissue. Additionally, incorporating the patients' own cells may facilitate the applicability of these types of constructs to contribute to both the structural and functional components of the tissue.