The use of 3D printing of artificial organs and tissues has expanded rapidly in the past decade and is expected to revolutionize the future of health care. Manufactured using an additive layer-by-layer method, these artificial constructs aim to replace the need for donor organs and can be tailored to the patient’s specific needs. The main goal behind this method is to create working three-dimensional vascularized organs and tissues that can be implanted into humans. This paper explores the vast field of 3D printing, including the materials, benefits, challenges, and risks associated with it.
3D printing is defined as the additive process of depositing a material layer by layer to create a structure . It has been widely used to create various prototypes — some examples include automobiles, toys, and computers . However, more recently, it has been adapted for use in the medicine. Rapid prototyping, of which 3D printing is included, has been used to generate models that could help with diagnosis, education, and surgery planning and preparation .
Before 3D printing technologies could be developed, computer-aided designing/drafting (CAD) and computer-aided manufacturing (CAM) technologies needed to advance. In 1956, what is considered to be the first CAM software, Automatically Programmed Tool (APT) was exhibited based on previous findings in numerical control and milling machines. The 1960s is when 3D CAD systems started making their mark in the industry, though there were still many drawbacks that kept them from becoming prominent in the industry. Early on, these technologies were separate entities — until the 1970s, when they started to be integrated with one another. By the 1980s, various modelers were available for CAD — including parametric modelers, feature-based modelers, and constraint-based modelers. In addition, the CAM technologies were able to interpret compound surfaces. Probably one of the biggest advancements in this field occurred in the 1990s: personal computers gained enough processing power to run these softwares, thus allowing the public to partake in this innovation. 
The first big advancement in 3D printing occurred in the 1980s: Charles W. Hull developed the process of stereolithography — which involves a liquid polymer that could be hardened with the use of a laser . Since then, 3D printing has been used for a variety of applications: rapid prototyping and manufacturing, replicating rare artifacts and fossils, customizing tools, and personalizing products . The 1990s marked a milestone in this technology, when the development of nanocomposites led to the ability to print more solid and reliable objects, thus leading scientists to consider the possibilities of 3D printing in the biological and medical fields — more specifically, 3D printing tissues and organs  .
While there are challenges that still need to be addressed, there is no doubt that the ability to print biocompatible materials, cells, and other supporting structures is a major milestone to utilizing 3D printing as a solution for fixing damaged or diseased tissues and organs .
The current strategy with tissue engineering is to directly deliver sets of biological “building blocks” into the body. These building blocks include cells, hormones, the extracellular matrix, and biodegradable scaffolds. The scaffolds can serve one of two purposes: as biological signals to prompt the body to repair itself (in vivo), or prior to the transplant, culturing these building blocks into more sophisticated tissues and organs (ex vivo). Bioprinting involves a spatially controlled deposition of biological inks to create two and three dimensional biomimetic patterns. 
Various materials can be used for 3D bioprinting, including naturally derived polymers, synthetic polymers, ceramics, composites, hydrogels, and metals   . The following cell types are used during 3D bioprinting: embryonic stem cells, adult stem cells, induced pluripotent stem cells, and tissue-specific cell lines .
Although transplantation is usually what is thought of when referring to 3D printed tissues and organs, there are other applications for these constructs. For instance, artificial organs could be used for drug discovery, modeling diseases, surgical preparation, and various testing and further research   .
There is a general process, seen in Figure 1, that needs to be followed when bioprinting. Due to the complicated anatomical structures of the human body, some form of medical imaging — such as X-ray, computed tomography (CT), or magnetic resonance imaging (MRI) scans — is usually utilized to generate a 3D image of the specific structure.  These images of the anatomic structure would need to be converted into a mathematical model that can be read by CAD software to create a 3D model design; followed by another conversion to present the data as STL files, so that the 3D printer can understand the shapes that need to be generated.  STL files contain various information for the printer to interpret, including shape, color, texture, and thickness of the object.  Essentially, an STL file represents the designed model in numerous triangular sections, like in Figure 2.  After generating a design of the organ, as well as a bioprinting process plan, the cells need to be separated and prepared to mature into the appropriate organ-specific cells.  Additionally, optimal materials need to be chosen to correspond with whatever structure is to be printed . At this point, the bioprinter recognizes the data provided as numerical controls, and the print head — or platform, depending on the method being used — will move in the x, y, and z-axes, accordingly .
In some cases, an additional step involving a bioreactor is incorporated into the process. After printing, the structure is placed into the bioreactor — this is the final step before transplantation.  It is important for the bioreactor to mimic the natural environment for which the 3D printed structure is meant. A bioreactor has the ability to produce necessary environmental cues, like physical and nutrients, to ensure that the structure matures correctly . As is the standard for any commercial product, quality assurance and various evaluations needs to be performed on the model to ensure that all needs and specifications are met .
As mentioned, CAD and CAM software is essential for 3D printing. Some examples of these technologies are AutoCAD, PTC Creo (Pro/ENGINEER), and MasterCAM.
PTC Creo, is a feature-rich and parametric modeler program developed by Parametric Technology Corporation. Since it is a feature-based program, instead of the user needing to design in accordance to low-level geometries — like faces, edges, and vertices — they have the ability to design through the use of features that carry engineering significance — such as tolerances, protrusions, coordinate systems, and assembly relationships. In addition, since the program is also considered a parametric modeler, which consists of the ability to assign constraints to the model, if the user modifies a single parameter, PTC Creo will automatically update any features that have a relationship with the modified feature, accordingly. This aspect of the program is especially useful when engineers need to make small but significant improvements on pre-existing products — as engineers frequently do. Instead of having to redesign the entire product, the old design can be used as a foundation for the new design, thus saving time and increasing efficiency. 3D boundary representation is incorporated in the program, and therefore, the designer is able to work with the part through the software before physically printing the prototype. This can help to anticipate, and therefore, reduce any design complications that may not have been originally accounted for. Since 3D printing a part could be approached in a top-down or bottom-up manner, PTC Creo has considered this and has made the platform able to accommodate for either approach. Furthermore, it provides engineers with other necessary tools: such as the ability to directly create engineering drawings, bills of materials, and parametric CNC tool paths — all of which can be connected to the original designed part. One of the better aspects of this software is that multiple people could work on a design at the same time while any changes made are updated in real time. 
AutoCAD has been available since the 1980s, and has been fairly successful. It allows for the easy conversion of design file data into data for CAM software to generate a NC tool path. Originally, AutoCAD was a 2D drafting software, but it has evolved into a 3D modeling software, as well. As a modeler, AutoCAD can generate designs using entities like extrusion, rotational sweep, chamfer, fillet, and more. Furthermore, the program can check for design feasibility and warn engineers about possible problems in the design. For example, AutoCAD can detect if parts in an assembly will overlap and cause an interference. However, unlike PTC Creo, the software does not allow for the integration of multiple users.
MasterCAM, which was developed by CNC Software, is a CAD/CAM software. This means that it is used for designing 2D profiles and 3D multi surfaces model of mechanical parts, as well as providing code to program automation of the machine tools for the manufacturing of the machine parts. An example of what a tool path, generated by MasterCAM, would look like is seen in Figure 4. This software is highly user friendly and interactive as it helps the user to edit the contour and surfaces of the machine part by selecting appropriate geometry once the design is made and varying the parameters; the output is then displayed in the visual format on the computer screen. It also contains features like solid-based tool path veriﬁcation which helps users see the process of cutting the solid block of material. The cutting time could be estimated by using the tool path back-plotting function. Another advantage of MasterCAM is that it can store a library of commonly used operations to automate machining and also provide a C-language interface — which helps in developing a customized program for specific applications. This tool can be used in 3D printing to design 3D Structures and edit them according to the need to make the design before actually starting the process. 
Types/Methods of Printing
There are various types of 3D deposition methods including, but not limited to, stereolithography, photolithography, selective laser sintering, fused deposition modeling, inkjet printing, pneumatic extrusion printing, electron beam melting, and laminated object manufacturing. Parameters that are important to consider when 3D printing include resolution, deposition speed, bioink compatibility, scalability, and its ease of use.  There is a strong potential to create more sophisticated artificial tissue constructs through 3D printing. Overall, biological printing gives researchers’ novel strategies to recreate constant microenvironments for both two and three dimensional designs.  This paper will focus on some of the printing techniques that are more commonly used in biomedical engineering.
Stereolithography could be considered the first 3D printing process, an apparatus of such seen in Figure 5, as mentioned earlier.  A laser emits an ultraviolet (UV) light to polymerize a photosensitive resin that is pooled in a container within the printing apparatus.  The light is cast in various shapes, in accordance to the image data provided, to account for the 3D structure that is to be generated.  Normally, stereolithography produces plastic parts; on the other hand, if the final product is desired to be consisting of metal, the plastic can be used as a pattern for the metal. 
Selective laser sintering, seen in Figure 6, uses a CO2 laser beam to fuse powder particles together, corresponding to a specified design. Once the product is completed, the surface needs to be sandblasted. While this method is highly accurate, it is also more expensive and requires a longer fabrication time.  The powder particles could be plastic, metal, or ceramic.  Factors that contribute to the resolution of the structure are the precision of the laser and the fineness of the powder used; it is possible to generate intricate and delicate products. 
Electron beam melting is similar to selective laser sintering. However, as its name suggests, it uses an electron beam (at about 4kW) instead of a laser to melt powder. Furthermore, the materials associated with this method are various metals, such as steel or titanium alloys. An example of a standard electron beam machine can be seen in Figure 7. 
In fused deposition modeling, a thermoplastic is heated as it is expelled from a nozzle, as seen in Figure 8. The common material used is acrylonitrile butadiene styrene (ABS) plastic — however, this is not a biocompatible material. The plastic is deposited in layers; over time, these layers will eventually produce a complete 3D structure.  The nozzle is able to move in both the y- and z-axis directions.  The system used for fused deposition modeling allows for control over the location of the plastic deposition. This technology is also less costly than that of selective laser sintering. 
There are several types of inkjet printing. A generic image outlining the inkjet printer apparatus can be seen in Figure 9.
One type uses the emission of an ultraviolet (UV) light to cure a photopolymer-based resin, much like stereolithography. However, in this case, the resin is ejected from a high-resolution inkjet and deposited onto a collection plate during the printing process. As layers build, a 3D structure eventually forms. 
Another type of inkjet printing is similar to selective laser sintering in the sense that powder particles are bound together. However, instead of a laser, some type of binding agent is used by being deposited from a print head. The surface is finished with cyanoacrylate-based material, which helps to harden the structure. Intricate shapes can be achieved at a relatively low cost and in a quick manner: the process, including printing and finishing, is estimated to take between 4–6 hours. 
The most common method of inkjet printing used is thermal inkjet printing, which involves using a power supply to heat the nozzle head, causing miniscule air bubbles to form that apply a pressure to the nozzle head, thus ejecting droplets of cells. 
Inkjets have a high level of compatibility with biomaterials, giving the user countless options to print with.  In each case, some type of “ink” is dropped from a printer head. The size of the drop that is jetted out can be manipulated by adjusting the temperature gradient, the pulse frequency, and the ink viscosity — the alteration in these parameters gives engineers the ability to produce a product with various mechanical properties. Importantly, it has been observed to have no adverse effect on mammalian cells.  Overall, inkjet printing is a process that is 100% programmable, which gives the user much leeway into printing the pattern the way it was intended. In addition, during the printing, there is no contact between the nozzle and the substrate, which greatly lessens the risk of contamination and allows the user to print on non-uniform substrates. Although the resolution of inkjet printing is lower than that of photolithography, it has been shown that the resolution generated by the inkjet printer is adequate enough for generate cellular responses in the printed construct. 
One of the major and most prevalent benefits of 3D bioprinting is the ability to make customizable structures. For transplantation, it is vital to consider the size of the organ being transplanted and the size of the body that the organ is being transplanted into. This can be an issue when the only organs that are available for transplant are too large for the patients that are in dire need of the organ. However, there are some instances when generally manufactured implants are preferable, but a custom-made implant proved to be more ideal. According to Aranda et al., a custom implant for a sternocostal reconstruction proved to have better functionality, fit, and cosmetics in comparison with non-custom implants.  A customized calcaneal prosthesis was found to be the optimal choice because the implant was able to be made hollow, thus reducing the weight.  Furthermore, in the case of the calcaneal prosthesis, since it was matched to the unique patient anatomically, it lowered the probability of the need for any allograft or autograft material adjustments in the future — which could be tedious for the patient to have to undergo — thereby, improving patient satisfaction . In both situations, the patients were discharged relatively quickly and rehabilitation was able to begin sooner.
Another benefit of 3D printing in medical applications are that items could be manufactured cheaply for small-scale productions.  While it is still considered more cost efficient to manufacture large amounts of products using traditional methods, many structures created for medical applications will not need to be generated on a large scale. For instance, implants would generally be uniquely produced on a case-by-case basis; therefore, they cannot be manufactured on a large scale. Moreover, the cost efficiency of 3D printing is beneficial when there are products that may need to be continuously modified, especially if those modifications only concern one part of a product.  Cost efficiency is also a result of eliminating the need for certain resources. Instead of designing an entirely new product, a 3D printer could be used to design only the part that needs to be replaced.
In addition, 3D printing is a generally fast technology, thus enhancing productivity. 3D printing is considered a faster approach than traditional manufacturing methods. Prosthetics and implants require various manufacturing methods including, but not limited to, milling and forging, which would take much longer than the hours it would take for these same items to be 3D printed.  For instance, the manufacturing time for a sternocostal implant was recorded to take about one month, while that of a calcaneal prosthesis only took several days.   These differing manufacturing times are most likely due to the following factors: the amount of parts printed for each implant, the complexity of the structures, and the overall size of the structures. Moreover, 3D printing improves resolution, accuracy, reliability, and repeatability. 
To develop a functional organ, one requires advances and integration of three types of technology: cell technology, biomanufacturing technology, and technologies for in vivo integration.  The main challenge is the adaptation of the existing system for particular biological material and to develop deposition techniques that are relevant to dispensing these particular materials. Also, there is a need to find an optimal way to put the technological components — the bioprinter, the container for dispensing cell aggregates, and the bioink together — this will contribute to the technology’s ease of use in a surgical setting.  It should be scalable and must be able to print a functional living organ that will be suitable for clinical implantation.  Another limitation with multicellular constructs is that the use of solid scaffolds can result in the low level of precision in cell placement. 
Viability of the cell during the printing process is one of the most important factors that needs to be kept in mind when designing a 3D printed organ. It is important to ensure that the method used for printing will not affect the survivability of the material.  Several photosensitive hydrogels have been used for fabrication to rapidly construct a desired geometry; however, it was seen that the viability and cell density was compromised since hydrogels restrict the movement of cells — thus, tissue cannot be formed.  To improve cell viability, researchers have developed cell friendly, biomimetic, photosensitive hydrogels containing arginine-glycine-aspartic acid (RGD) peptide specific for cellular integrins.  Another solution to maintaining cell viability could be to develop novel irrigation perfusion bioreactors based on using temporal, removable, porous needles with pressure controlled, dripper-like systems. 
As mentioned, there are various materials that have been used for bioprinting. These materials correspond to one of the major challenges in bioprinting. The material should be biocompatible and have the same mechanical and functional properties as the native organs.  The materials should be selected with consideration of the following properties: material chemistry, molecular weight, solubility, shape and structure, hydrophilicity/hydrophobicity, lubricity, surface energy, water absorption degradation, and erosion mechanism.  It should not induce any undesirable systemic responses when it’s implanted into the body.  Depending on the application, materials can be synthetic or natural, degradable or non-degradable. Naturally occurring polymers are preferred because of their bioactive properties; synthetic polymers are cheaper than biological scaffolds, and have a long shelf life. However, synthetic polymers have poor biocompatibility, toxic degradation byproducts and decreased mechanical properties during degradation, which restrict their use.  The optimal method is to use a blend of synthetic and natural polymers that improve the scaffold properties, thus allowing controlled degradation. Over time, a scaffold degrades, which introduces a risk that, for a given tissue, it may not match the ability of the cellular components to replace the material upon degradation.  Thus, the necessary precautions should be taken to ensure that the degradation of bioproducts are nontoxic, readily metabolized, and rapidly cleared from the body so that it does not have any harmful effect. 
Bioreactors play a very important role in the development of 3D printed organ. Bioreactors provide cells and molecules with the appropriate environmental cues for them to differentiate and proliferate appropriately — therefore, playing a vital role in the production of a functional tissue or organ.  Bioreactors should be designed in such a way that it should allow easy removal of the printed tissue construct without damaging it.  The bioreactor should have specific characteristics, such as allowing perfusion of the intra organ branched vascular tree, providing a timely removable irrigation system to give enough time for the vascular system to mature and start intravascular perfusion. During post processing, it must give dynamic biomechanical conditioning that will help speed up the tissue maturation. Accelerated tissue maturation is a notable challenge due to the fluidic nature of manufacturing and the absence of a solid scaffold. In order to get good mechanical strength, the printed organ construct should be fused, remodeled, and matured quickly as proteins, such as collagen and elastin, are produced in abundance during maturation. The transfer process of printed organs to the bioreactor should be done carefully. There should be no vibrations as it may compromise the mechanical integrity of fragile gel-based structures. Most importantly, it must integrate with the bioprinters and other components without damaging the tissue. 
Another important step in manufacturing 3D printed tissues is to observe the kinetics during remodeling and maturation. Today, biosensors and other quantitative methods are used; however, it is not foolproof. Thus, there is a need to develop noninvasive, nondestructive methods of monitoring fluid flow during maturation. Moreover, another challenge to bioprinting is to reproduce the complex microarchitecture of the extracellular matrix component. A sufficient resolution that resembles the original biological function is required. However, it is difficult to mimic the cellular organization of natural tissue, and it is an expensive technique to reproduce.   
As promising as 3D printing is for medical applications, the ultimate challenge will be achieving FDA clearance. Even if the results of various medical studies are promising, it must first pass through the United States Food and Drug Administration. As more biological inks and constructs surface, the amount of time needed for them to get cleared by the FDA also increases. Once passed through the FDA, the next step would be to show the product’s competitiveness in its respective market. The rate of adaptation among clinicians will heavily rely on its cost effectiveness compared to other existing solutions such as allografts. 
As with any medical institution, there are ethical principles which orders what the community will regulate concerning the human body. The ethics and risk behind 3D printing artificial organs and implanting them into humans and animals have become a debated topic. Specifically, health care issues have come into play concerning the price of these implants. If 3D printed organs become mainstream, the price is something that will be an issue with patients. The question arises of at what point is the option of having an artificial organ be denied to a patient in need? The solution would be for the FDA and health insurance companies to set regulations on how frequently organs and tissues are donated and replaced, however such a collaboration will take years to be fully implemented.
The Biomedical Engineering Society Code of Ethics Obligations states that Biomedical Engineers must “use their knowledge, skills, and abilities to enhance the safety, health, and welfare of the public”. An example to consider is with artificial heart valves. If one is implanted into a patient, the clinician must factor in the risk of having to replace it again after a few years. Depending on the health of the patient, multiple surgeries might not be an option. A way this could comply with the Biomedical Engineering Code of Ethics is to devise a heart valve that has the ability to conform in a young patient and grow as the patient grows older. However, to achieve this, much clinical testing is needed on human subjects, which puts them at risk, and therefore goes against ethics. It comes down to which problem holds a greater violation: denying a patient a potentially life-saving artificial implant, or going through with it and risking their life to using it. 
Risks and Possible Design Errors
After developing any novel technology, it is always subjected to patent, copyright, and trademark law. However, for 3D printing technology there has been very little exposure on how these laws are to be applied for commercial and personal use. In order to sell 3D-printed items that are patented, the person has to negotiate with the owner of said patent to get licenses so that he or she can sell the product. Copyright is another issue; the designer who files for copyright always demands for removal of the design of other participants from the repository because those designs may encroach on the copyright. Meeting all the FDA regulatory requirements and getting a final approval is one of the main barriers that hampers the widespread use of medical applications of 3D printing on a large scale. 
Another significant risk in organ printing is the integration of a vascular network. However, the conventional tissue engineering approach is not capable of creating complex vascularized organs. Instead, two main strategies were proposed for inducing vascularization of tissue-engineered organs: incorporation of growth factors into the scaffold to induce angiogenesis after implantation or the pre-seeding of the implant with endothelial cells.
Unfortunately, both of these strategies are not foolproof as there are complications that arise due to the slow rate of vascular tissue remodeling and the necessary highly branched vascular networks for maintaining viable cell constructs.  Usually, the organs that are made are small and relatively simple. They are avascular, aneural, alymphatic, thin, hollow, and are nourished by the diffusion from the host vasculature. However, if the thickness of the engineered tissue exceeds 150–200 micrometers, this causes problems in oxygen diffusion between the host and transplanted tissue. Large organs, like the heart, kidney, and liver cannot maintain metabolic functions without vascularization.  Thus, functional vasculature needs to be bioprinted so that there is a constant supply of oxygen, nutrients, growth factors, and also the removal of waste products, which is necessary for the maturation and growth of cells during perfusion. Usually, this results in low cell viability and malfunction of the artificial organs and therefore prevents organs from growing properly.  Thus, the system must be developed to transport nutrients, growth factors, and oxygen to cells while extracting metabolic waste products such as lactic acid, carbon dioxide, and hydrogen ions so the cells can grow and fuse together, forming the organ. 
Biological ink is an important component in 3-D printing of artificial organs. It is responsible for encouraging desired cellular interactions to ensure the printed cells can proliferate and differentiate in a scaffold. Biological inks are composed of active mechanisms such as, but not limited to, proteins, DNA, growth factors, extracellular matrix molecules (such as collagen), natural and synthetic polymers.  The main parameters to working with bioinks are to control the physical and molecular properties of the extracellular matrix (ECM). By controlling the behavior of the cells, there is a greater control over how the tissue integrates with the scaffold. A popular way to do this is through the use of natural polymers, due to its similar components to natural ECM. Among the choices for natural polymers are dextran, collagen, gelatin, alginate, fibrin, and hyaluronan. Another option is the use of synthetic polymers. An advantage to these polymers is the greater flexibility in creating a polymer that suits the needs of the user. Some examples of synthetic polymers being used today are poly(ethylene glycol) (PEG), poly(L-Lactic acid) (PLLA), poly(ɛ-caprolactone) (PCL), and poly(lactide-co-glycolide) (PLGA). However, when considering the biomaterials for 3-D printing, researchers are not limited to just using either synthetic or natural polymers. 
Combinations of the two have been studied to bring out the advantages of both types. When designing biological inks, there are many parameters to consider. The properties of the material chosen must be meticulously examined to ensure no problems will occur during the printing process. The parameters that are explored include viscosity, hydration, biological interaction capabilities, mechanical strength, polymer chain properties, and techniques of fabrication. 
It can be said that one of the most important parts of the printing process is the precise moment when the bioink is injected from the nozzle. An ideal polymer will start off as a fluid inside the printing chamber, and end as a solid once it is deposited onto the desired surface. Viscosity of the polymer decides whether or not it will smoothly eject from the tip onto the scaffold. If the viscosity is too high, there is a risk of clogging, and if it is too low, the shear rate will kill the cells before they have a chance to make it to the scaffold. One downside to using natural polymers is the variation in quality between batches of bioinks. The batches must have constant properties to ensure a balanced, cohesive structure. Thus, it is important to contain the bioink in a controlled environment that will prevent it from being adulterated. Using gels to surround the ink will protect its structure better than leaving it in a fluid state. Hydrogels can be produced using a semi-interpenetrating polymer network. This can account for the gel’s mechanical strength and hydration due to a secondary network of uncross-linked polymer. Two sub factors that can affect viscosity and the overall strength of a bioink are hydration and porosity. In order to keep the printed cells alive, there needs to be a supply chain of oxygen and other nutrients within the hydrogel. Both natural and synthetic polymers possess the capabilities to act as natural tissue due to its hydration properties and ease in forming structures. A disadvantage to this is the higher pore size, which is linked to a lower mechanical strength. The greater the pore size, the more contact there is between water and the hydrogel. Water is seen to function as a plasticizer on the hydrogel, which can compromise the structure if there are thermal changes applied to it.     
In order to mimic the natural cellular environment, the mechanical properties of these tissues must be understood. Natural tissue is made up of a hard acellular matrix that is encircled by a soft cellular matrix. Temperature is another physical factor that has to be regulated during the printing process, as polymers are sensitive to thermal change. When looking at polymer chains, one must consider how the presence of water can affect the material. It can alter the mechanical stability of the biomaterial, and can affect the supply chain of nutrients and cycle of metabolic waste within the polymer. Degradation is an essential factor to consider when planning a printed structure. Polymers susceptible to hydrolysis, such as polyesters, influence different degradation rates in structure made from synthetic polymers. Natural polymers contain sites where cleavage can occur through the use of enzymes, which is useful for cell migration through a hydrogel. This is because the use of enzymes triggers the biosignals in the surrounding matrix to decrease its mass either externally or internally. 
The polymer is essentially the “make or break” factor in any bioprinting process. The advantages to using a natural polymer is its compatibility in a simulated extracellular matrix. It is capable of composing constructs of various parameters in this manner. There are multiple natural polymers in use today.
The first is Alginate, which is a polysaccharide derived from seaweed. It is known for its physical similarities to the extracellular matrix, its biocompatibility, and its ideal level of viscosity and level of gelation at room temperature. A drawback to using Alginate is its lack of cell-binding functional groups, which could inhibit adhesion and proliferation of the seeded cells. In the case of cell adhesion, this can be achieved by combining Alginate with gelatin, since it contains the necessary adhesive cells. However, there are some remaining problems using alginate as a bioink. It is difficult to achieve optimal porosity levels in alginate-based scaffolds due to the difficulty in handling the gelation of this polymer. Another problem which can arise is the medium into which is it deposited into. Alginate has a high level of solubility in aqueous media, which can prevent the printing of thicker structures. An alternate solution to this issue was proposed by Ahn et al. It involved using a different printing process consisting of depositing the alginate and then using an aerosol-spraying procedure to coat it with CaCl2. After which, the scaffold was submerged in CaCl2. The process was altered by imbedding bioactive “clues” such as proteins into the scaffolds before spraying the CaCl2. For example, bone morphogenic protein 2 (BMP-2)-loaded gelatin microparticles were placed in the scaffold and displayed osteogenic properties when evaluated in vivo. In doing so, the scaffold had a much more controlled release of the protein. Thus, researchers are hopeful that by using these bioactive clues, the processes in alginate based scaffolds can be controlled and be made a more viable choice for bioinks.   
A major component of the extracellular matrix is collagen. It is present in many physiological processes and influences cell adhesion, differentiation, proliferation, and migration. As a result, it is one of the more popular choice when creating scaffolds, especially ones based for tissues. Collagen itself is composed mostly hydroxyproline, glycine, and proline. In addition, it contains trace amounts of amino acids containing sulfur and other aromatic compounds. When it undergoes hydrolysis, collagen gives off gelatin as a byproduct. Gelatin is a triple helix structured polypeptide derivative. It is nontoxic, possesses biocompatibility, biodegradable, and can influence adhesion. Like collagen, gelatin is also a prime choice for bioinks due to its similarity to collagen. 
An added benefit to using gelatin is because of its denaturalized state when derived from collagen. This state inhibits the spread of pathogens and any risk of immunogenicity. However, gelatin is a temperature dependent polymer. In some cases, where temperature cannot be tailored for the gelatin, it is usually combined with another polymer, such as fibrin. Other methods included further chemically altering the gelatin, such as having it undergo methacrylation and acetylation. Acetylation helped jump start the cellular flow in the gelatin mixture. Exposing the gelatin to methacrylic anhydride helped identify the mechanical properties of the mixture. Using high amounts of it (around a 10 molar excess) resulted in low levels of viscosity of the gelatin. However, lower amounts of it resulted in soft hydrogels, but with a higher level of viscosity. This parameter was controlled by further acetylating the gelatin.  
Chitosan is a porous derivative of chitin that has enhanced liquid absorbing capacity and cell interaction. It is biocompatible, biodegradable, nontoxic, and possesses antibacterial capabilities. Its most significant drawback is its brittleness When cooled to lower temperature, a portion of Chitosan remains uncrystallized, which means that it will not dissolve in an environment that has a pH of over 7. This property makes chitosan an option for bioprinting. The chitosan solution by itself is viscous. To deposit it and induce gelation from acetic acid, sodium hydroxide is used. A NaOH concentration of 0.75%-1.5% v/v was optimal for use with Chitosan. Concentrations above the optimal range resulted in faster rates of gelation, whereas lower concentrations caused the gel to spread over the printed struts, which is undesirable. The excess NaOH can be removed after the printing process by immersing it in ethanol and then stored in deionized water. Scaffolds composed of Chitosan displayed well proliferation of seeded cells.  
Another major component of the extracellular matrix is Fibrin. It is created from the interaction between the glycoprotein Fibrinogen and the serine protease Thrombin. Fibrinogen is composed of several pairs of polypeptide chains such as Aα, Bβ, and γ. The thrombin cleaves the Aα and Bβ chains to produce fibrinopeptide A and B, which directly causes a chain reaction in cellular changes. The fibrinopeptide polymerizes into protofibrils, to fibrin fibers, and finally into fibrin gel through association. The reason Fibrin is potentially viable for use in bioinks is its adhesion cell signals found in Fibrinogen. There were favorable results when looking at how seeded cells proliferated on scaffolds composed of fibrin and the level of cytocompatibility displayed. However, there are some drawbacks to using fibrin-based bioinks. In its gel form, fibrin is not as elastic as other bioinks, rather it is fragile. An element that is essential for bioinks is its ability to maintain its printed shape, something which is not found in fibrin. The drawbacks of using bioinks composed of only one polymer can often be solved by using composite bioinks.   
Using Fibrin as an example, it has been used in conjunction with collagen as a bioink. Since both are major components in the extracellular matrix, it can be counted on to have a strong biocompatibility. An example of its use is in treating skin wounds. The Fibrin-Collagen bioink, combined with amniotic stems cells and bone marrow mesenchymal stem cells were seeded onto wounds. Due to the morphology of the Fibrin and Collagen (rod-like shapes), the proliferation of the cells was restricted to an extent. In addition, the signals for cell adhesion found in Fibrin can also play a part in restricting its proliferation, which is a drawback for the healing process.  
As mentioned before, a drawback to using Alginate is its lack of cell-binding functional groups, which could inhibit adhesion and proliferation of the seeded cells. Cell adhesion can be achieved by combining Alginate with gelatin, since it contains the necessary adhesive cells. Gelatin ensures that the bioink’s flow rate is tailored in such a way that the deposition from the nozzle ensures it is stable upon printing before being cross-linked with the Alginate. Another upside to using Gelatin is its thermoresponsive traits. By cooling the Gelatin-Alginate bioink to under 10°C, it transforms to its gel-state, which allows for the cross-linking to occur without compromising its integrity When using such a composite bioink, it is important to consider the ratio of Gelatin to Alginate. Having more Alginate resulted in decreased mechanical strength of the deposited struts. On the other hand, a higher ratio of Gelatin caused the bioink to be more viscous, which can cause clogging to occur in the nozzle when depositing the bioink. Thus, a balanced ratio of the two is necessary to insure smooth deposition and viability in the long run.  
Another polymer that can be a practical bioink, but limited if used by itself is Hyaluronic Acid. Found in the extracellular matrix, it is a polysaccharide that is biocompatible, biodegradable, and viscoelastic. However, it has a highly water soluble, which can compromise its stability if printed in its natural form. There were several attempts to use Hyaluronic Acid with another polymer in order to mitigate this drawback. Pescosolido et al. was able to do this the best by combining Hyaluronic Acid with dextran, specifically a derivative of it: hydroxyethyl methacrylate derivatized dextran. It is used as a hydrogel by radical polymerization of the derivatized dextran with a reactive group such as esters. It is then stabilized using UV irradiation The mechanical strength of the Hyaluronic Acid and Dextran bioink can be controlled by changing the concentration of the dextran derivative in the blend. With the dextran derivative itself, changing the degrees of its substitution also has an effect on the mechanical strength. In addition to its ease of use, at a low shear rate, the composite is very viscous, making it an even more viable option for bioprinting.  
Although natural polymers have the ability to mimic the extracellular matrix, it often comes with drawbacks in terms of mechanical and chemical properties. The use of synthetic polymers allows the printed constructs to have enhanced mechanical strength, biocompatibility, and better rate of degradation. When combined with molecular agents, these synthetic polymers can be used to generate responses through the mechanotransduction pathways. Some of the most commonly used synthetic polymers include Poly(ethylene glycol), Poly(lactide-co-glycolide), Poly(ε-caprolactone), and Poly(l-Lactic Acid). 
Poly(ethylene glycol), or PEG, is a synthetic polymer that is widely used in biomedical engineering as a material for tissue engineering, drug delivery, and bioinstrumentation due to its hydrophilic and biocompatible nature. Like Hyaluronic acid, it is water soluble, making it unable to form a hydrogel in its natural form. To use it as a bioink, PEG is acrylated by crosslinking its chains through polymerization under UV irradiation. These hydrogels have been shown to have high mechanical strength and being able to maintain its form under pressure. 
Poly(lactide-co-glycolide), or PLGA, is composed of two monomers: lactide and glycolide, created by the process of ring-opening polymerization. It has garnered much attention as a material which has controlled degradation rates and strong mechanical properties. When working with a multifaceted structure which requires circulatory networks, flow control, etc., PLGA enters as an option to satisfy these requirements. 
The Ringeisen group though of the idea to use PLGA as the “building blocks” for cardiovascular cells to generate high resolution three dimensional tissue through two dimensional biological laser printing. An example of an experiment utilizing PLGA involves dissolving it in chloroform and placing it into Polydimethylsiloxane molds with salt. After washing out the salt and letting it dry, the created scaffolds had either collagen or collagen and Matrigel inserted, which served as the biological media. Onto the scaffolds, human umbilical vein endothelial cells were deposited in the “building block” manner mentioned. This displays PLGA’s element of allowing the proliferation of two dimensional cell patterns to form a three dimensional construct. Thus, as an additive, PLGA is an excellent biomaterial to use in conjunction with growth factors. PLGA’s ability to degrade in the presence of water ensures that during the regeneration process, it will not hinder the cells to proliferate and construct into its intended structure. 
Poly(ε-caprolactone), or PCL, is another semi crystalline polymer, used in conjuction with other polymers to overcome the brittleness found in PLA and plastics. Like PEG, PCL is used in tissue engineering due to its thermoplastic nature, notable mechanical strength, being biodegradable, and possessing a low melting point (reported around 60°C), which adds to its ease of use. Like PLGA, PCL is used as an additive because its melting point is too high for cell survival. Therefore, it needs to be used with another bioink, usually as a means of support for it. Unlike natural polymers, synthetic polymers such as PCL do not mimic the extracellular matrix, rather it supports the printed construct. This category is of systems are known to be “scaffold-free”, which rely on the cells ability to form into its own structure through cell-cell interactions.  
Poly(l-Lactic Acid), or PLLA is semi crystalline polymer that is both biocompatible and degradable. It is known for its ease of use in deposition, as it displays a low viscosity from the nozzle, which allows for a smoother flow rate. After it is ejected, it evaporates quickly to form a firm construct. PLLA was used by the Therriault group to fabricate various microchannels using a ”solvent-cast direct-write” method. It was shown that PLLA could fabricate different morphological constructs in the shapes of square and circular spirals using layer-by-layer deposition. Similar to the Hyaluronic Acid-Dextran composite, PLLA has a low shear rate, allowing the ink to deposit efficiently. At the atomic level, PLLA has a greater level of mechanical strength over other synthetic polymers such as PCL. PLLA has a shorter alkane chain in its backbone, and possesses more methyl pendant groups, which inhibits the backbone of the PLLA to rotate.  
As mentioned before, the use of composite natural polymers helped alleviate the physical drawbacks of using uniform bioinks, however, the problem remained of trying to control the resulting biological processes. Thus, researchers have tried combining natural and synthetic polymers to draw out the advantages in both. For example, synthetic polymers have been shown to a useful additive in providing structural support. It can be combined with a natural polymer to incorporate its biomimetic properties and create better adhesion and proliferation. A problem to using this type of mixture arises during the printing process. Natural and synthetic polymers are not chemically bound to each other, which can affect the overall viscosity of the composite bioink (specifically, natural polymers tend to be more viscous for more control during deposition). Thus, instead of using one bioink composed of natural and synthetic polymers, a different approach is taken: designing a hybrid ink system consisting of a synthetic polymer ink to create the scaffold, and a natural polymer ink as the “packaging” for it. There are numerous benefits to using this hybrid system. Synthetic polymers, such as PCL, provide the mechanical backbone for the scaffold and a sufficient rate of biodegradability, while allowing more multiple hydrogels to be deposited onto it. The strength provided by the synthetic polymer allows for less viscous hydrogels to be printed onto it, which augments the number of bioink materials that can be used in conjunction.    
Metals are also possible materials to use in 3D printing. Titanium, in particular a surgical grade titanium alloy, has proven to be a feasible choice for applications, such as implantation, in the past due to its biocompatibility . 3D printing, which can be done through electron beam melting, this metal for use in medical applications, due to its strength and flexibility, has been an ideal option for customization.   Furthermore, titanium is light — thus reducing weight, however expensive. 
This section will look at some of the current 3D-bioprinters currently being used in the market. Many of these printers for sale consist of a pressure-based extrusion of both polymeric substances and hydrogels (acting as the bioink) through a syringe.
1) EnvisionTEC’s 3D Bioplotter
EnvisionTEC’s 3D bioprinter consists of a syringe-based extrusion system that can utilize biomaterials such as chitosan and hydroxyapatite, along with hydrogels. 3D models and patient data are taken and converted into a three dimensional scaffold. Thus, it is a useful research tool to experiment with different biological “inks” and their subsequent effect on the scaffold. The price for this printer exceeds $200,000. 
2) Organovo’s NovoGen MMX
Organovo is a company located in San Diego, California. They are one of the leaders in bioprinting and tissue engineering. The NovoGen MMX operates by taking cells and growing it in a culture until a sufficient number of cells are available. It is then harvested to create a bioink and then be loaded. Using computer scripts, it then deposits the tissue material into its target shape. The unit is currently not for sale. Organovo uses it to produce these specialized tissue to be then sold to major pharmaceutical manufactures for drug testing.  
3) RegenHU’s 3DDiscovery + Biofactory
RegenHU is a company based in Switzerland that focuses on creating commercial 3D bioprinters. The product consists of two components that work to combine various ECMs, hydrogels, polymeric materials, etc. within a sterile environment. In addition to hardware, RegenHU also markets their own line of bioinks, most notably the OsteoInk, which is a calcium phosphate ink that is ready to use for bone composition. The price of this printer exceeds $200,000.  
4) BioBots BioBot1
Based on the research conducted, there is no question that there are many improvements that need to be made on 3D printing technology to make it feasible for certain medical applications: like transplantation or implantation. It is our personal opinion that, at its current state, 3D printed artificial organs and tissues would be best used for the function of generating models as visual aids for surgeons and patients. This utilization would not be limited by challenges, such as the need for vascularization or biocompatible materials. This functionality is already implemented in clinical trials applications. However, we also believe that research towards the use of these structures for transplants or implants should continue due to the benefits 3D artificial organs and tissues have the potential of providing.
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Originally published at medium.com on February 12, 2017.