Additive manufacturing for biofabricated medical device applications

Additive manufacturing for biofabricated medical device applications

9

Michael P. Francis*,**, Nathan Kemper**, Yas Maghdouri-White**, Nick Thayer** *Eastern Virginia Medical School, Norfolk, VA, United States; **Embody LLC, Norfolk, VA, United States

Chapter Outline 1 Background 312 1.1 Introduction 312 1.2  Medical devices  312 1.3 Challenges 315 1.4  Opportunities for advanced additive manufacturing in medicine  316 1.5 Synopsis 317

2 Advanced biological materials in additive manufacturing  317 2.1 Introduction 317 2.2  3D scaffold printing of biopolymers  318 2.3  Synthetic polymer printing  319 2.4  Biological polymer printing  320 2.5  Cell printing  321

3 Advanced biomanufacturing modalities for medical devices  323 3.1 Electrospinning 323 3.2 Fabrics 325 3.3 Molding 326 3.4 Microfluidics 328 3.5  Pop-Up Book microelectromechanical systems  328 3.6  Soft robots  329 3.7 Metals 330

4 Production challenges  330 4.1 Introduction 330 4.2  Sterility and packaging  330 4.3  Manufacturing challenges  333

5 Summary 334 References  336

Additive Manufacturing. http://dx.doi.org/10.1016/B978-0-12-812155-9.00009-8 Copyright © 2018 Elsevier Inc. All rights reserved.

312

Additive Manufacturing

1 Background 1.1 Introduction Additive manufacturing (AM) for advanced medical device applications (biofabrication) encompasses a versatile, growing array of technologies for generating new twodimensional (2D) and three-dimensional (3D) materials, synthesized layer-by-layer, for both living (e.g. cellularized scaffolds) and supporting devices (e.g. scaffolds, mechanical supports, or prosthetic devices). In traditional manufacturing (TM) methods of medical devices, feedstock material is removed from a bulk substance (e.g. titanium bar stock or allograft bone that is cut via a lathe or by computer numerical control [CNC] mill) to form the therapeutic implant materials or support devices. In contrast, AM deposits, fuses or otherwise builds layers of material anew. AM, particularly for biologics, typically builds materials in small quantities and at relatively slow speeds, such as by layering endothelial cells oriented in a circle upon many layers of polymer to form a 3D blood vessel analog [1]. AM allows opportunities for precision manufacturing of materials that otherwise are not readily manufactured, or even attainable, by TM methods. A comprehensive features list (Table 9.1) of traditional and biological AM techniques as well as diagrams (Fig. 9.1) of each is shown below.

1.2  Medical devices This chapter focuses on AM applications for medical device products. The United States Food and Drug Administration (FDA) defines a medical device as “an instrument, apparatus, implement, machine, contrivance, implant, in vitro reagent, or other similar or related article, including a component part, or accessory which is: Recognized in the official National Formulary, or the United States Pharmacopoeia, or any supplement to them, • Intended for use in the diagnosis of disease or other conditions, or in the cure, mitigation, treatment, or prevention of disease, in man or other animals, or • Intended to affect the structure or any function of the body of man or other animals, and which does not achieve any of its primary intended purposes through chemical action within or on the body of man or other animals and which is not dependent upon being metabolized for the achievement of any of its primary intended purposes.” •

This definition of medical devices covers a wide array of products, and with unique sets of risks and benefits. Devices are further classified by the FDA into three main categories: A Class I device is of low risk (e.g. dental floss); Class II devices pose higher risk with more extensive regulatory control to assure safety and effectiveness (e.g. a woven polylactic acid scaffold for rotator cuff reinforcement); and Class III devices, with the highest level of associated risk and regulatory control (e.g. an artificial heart valve). Class III devices require premarket approval (PMA) by the FDA before marketing. While some devices are exempt from approval, such as human cell and tissue products (HCT/P), most Class I and II devices, particularly implantable devices, required a review via the 510(k) process. Presubmission meetings with the FDA are recommended for all new devices to discuss study plans and critical test plans for data acquisition to

Table 9.1  Summary

of characteristics of common AM modalities

VAT photopolymerization

Material jetting

Binder jetting

Med

Low

Print Speed Resolution (min, µm) Material types

Med-High 25–300 Photo-polymers and plastics

Cost Pros

Low Low-High Low-High Affordable, Very high speed Very high highand accurate, speed, speed and commercially multiple commercially available, colors, available multimaterial many builds materials Limited materials Limited Binder Low speeds, and high materials reduces part high cost maintenance strength, postprinting steps add time

Prep time

Cons

ABS, Acrylonitrile butadiene styrene.

Bio-extrusion

Fusion deposition

Powder bed fusion

Direct energy deposition

Laser-assisted bioprinting

Low-Med

Low-Med

Low

Low-Med

Low

Med-High

High 25–300

High 25–300

Low-Med 5–100

Low-Med 15–300

Low-Med 20–200

Low-Med 1000–4000

Med-High 5–500

PP, HDPE, PS, PMMA, PC, ABS, HIPS, EDP

Stainless steel, ABS, PA, PC, glass

Plastics, Thermoplastics, Metals, ceramic Stainless Hydrogel, polymers, wax, powders, steel, media, cells, hydrogels, elastomers thermoplastic cobalt proteins and media, cells, polymers chrome, ceramics proteins titanium Med-High Low-High Med Med-High High Designed for Very Wide range of Material High speed biocompatible inexpensive materials, grain printing for materials, can and no support structure biomaterials keep cells commercially material tunable to viable available needed application Part quality lower, accuracy and speed lower than other methods

Parts lack structural integrity, low speeds

Limited to metals, rough finish

Long prep time, limited materials

314

Additive Manufacturing

Figure 9.1  Diagram of commonly used additive manufacturing modalities.

support either a 510(k) or PMA application. The FDA Product Classification Database (https://www.accessdata.fda.gov/scripts/cdrh/cfdocs/cfpcd/classification.cfm) should be referred to for any new device classification, and to determine if any exemptions exist. A 510(k) pathway may be applicable to AM medical devices if the product is substantially equivalent to an approved medical device in terms of composition, specifications, and instructions for use, so long as special controls are put in place by the manufacture to assure the quality of the product. As a simplified regulatory case, a new tendon antiadhesion medical device manufactured via extrusion to 3D print a sheet of polytetrafluoroethylene (PTFE) may be cleared by the FDA as a 510(k) Class II device even if the AM method employed is unique. If the AM PTFE antiadhesion

Additive manufacturing for biofabricated medical device applications

315

device has comparable performance to the so-called predicate device made of PTFE manufactured by TM methods, the new AM product, if used in the same indication, with identical instructions for use, would likely be classified the same. The new AM device would have to further meet the data requirements set for by the FDA, adhere to special labeling requirements, have postmarket surveillance performed, and follow the general guidelines set for by the FDA during the approval process. Where the 510(k) path may be viable for AM of less complex biomaterial scaffolds alone or for diagnostic tools, AM involving living cells or growth factors (i.e. drugs), such as bioprinted living grafts that are intended for implantation for any indication, would likely require PMA. Further, for such an AM living or drug tissue, an investigational new drug notification must be sent to the FDA with a summary of the investigational plan to evaluate the drug or biologic. For this reason, many investigators and companies have favored the testing and commercialization of AM products without biologics (i.e. cells or small molecules, growth factors, or other drugs) to speed time and reduce costs for getting into the market with a first-generation product. After establishing a platform for the new AM technology, a next-generation product with added biologics becomes more reasonable for many indications, given the risks and added complexities involved. For data collection before FDA review process, standard testing is often required (e.g. ASTM and ISO). Nonclinical laboratory studies that will likely be required for most AM medical devices to include information on microbiology, toxicology, immunology, biocompatibility, wear, shelf-life, and other bench or animal testing. These tests must comply with Good Laboratory Practices (21CRF Part 58). Clinical investigations will additionally require study protocols, safety and efficacy data, complications and adverse reactions, recording of device failures and replacements, patient information and complaints, statistical analyses, and any clinical results to be gathered and approved prior to device marketing.

1.3 Challenges Akin to the challenges of AM in the fields of energy, aerospace and motorized vehicles, major societal advancements in AM for medical devices come at the cost of a high level of design and technological complexity, in a highly regulated and standardized environment. Advanced AM technologies require highly skilled workers to operate and maintain complicated equipment. Envisioned bedside AM of tissues and organs [2] in the future would present incredible hardware challenges beyond clean room manufacturing of a typical medical device. Critically involved in biomedical AM strategies are high performance computing and computer technologies, particularly those involving 3D rendering (computeraided design, manufacturing, and engineering [CAD, CAM, and CAE]). To accelerate manufacturing and enable precision or personalized device manufacturing, high performance computing technologies for AM often need to further be integrated into robotic and medical imaging systems, such as magnetic resonance imaging (MRI) and computed tomography (CT). As a human liver contains around 3.61 × 1011 cells, with varied cells of at least a dozen major types, the computing power and hardware ability

316

Additive Manufacturing

for manufacturing a human organ of this complexity is beyond current capabilities. A relatively acellular tissue such as bone contains 98.97 ± 1.24 cells/mm3 in trabecular bone and 56.34 ± 1.69 cells/mm3 in cortical bone [3], making AM of bone or other hypocellular or small volume tissues or organs (e.g. ligaments, tendons, retina) more promising initial targets, where viable products are the goal. AM technologies also require extended research and development time and cost to bring to maturity. AM challenges include the need for precision tooling required for AM equipment, manipulation of nanomaterial or nanofiber, control of manufacturing particulates in a clean manufacturing environment, and even sustaining of life in cellularized devices, among other challenges.

1.4  Opportunities for advanced additive manufacturing in medicine AM approaches in medicine have potential for generating intricate arrangements of cells, engineered scaffolds, tissues, and organoids which cannot be produced by other methods. The ability to control the spatial relationships of cells, matrix, and even growth factors (e.g. morphogenes, the factors which guide embryonic development) in 3D space may lead to advanced regenerative and replacement therapeutic medical devices, and to developing diagnostic tools. Current biomaterials used in 3D bioprinting often utilize polymers for support (e.g. polycaprolactone (PCL) or PLA) or hydrogels (e.g. gelatin or Matrigel®) to generate substrates within or upon which cells can be plotted, and to which growth factors can be tethered. The current common biomaterials, while effective for pilot research work, cannot be translated to clinical use, which is a major concern for a product-centric technology. For example, current 3D bioprinting approaches often produce mechanically unsound gels where solid structures are needed therapeutically [4], or use polymers which lack biological enhancement and performance comparable to a native biological matrix in terms of cell-matrix and cell-growth factor-matrix interactions [5]. Most existing materials used in AM are also limited simply by not being familiar or previously approved by the FDA, complicating the commercialization strategy of a 3D bioprinted material or tool. For scalable device manufacturing of 3D bioprinted materials, the generation of advanced biomaterial filaments beyond those that are commonly available will provide significant advancements to the field. Trends in the field are moving toward generating filaments with native extracellular matrix (ECM) cell-binding motifs and have mechanical properties that match with the native tissue biomechanics. These materials present cues to cells and the tissue surrounding the implant, which are not present in current printed synthetic polymer alone or as hydrogel constructs. Generating such materials requires advanced biochemistry, ECM biology, polymer science, and chemical engineering approaches, and therefore requires significant interdisciplinary collaborations, and extraordinary cross-functional team management for success. Ultimately, the development of new materials and the refinement of current AM filaments are needed to enable production of robust, reproducible scaffolds that are tailored to specific medical indications. Manufacturing of these materials could include new AM methods, improved polymer synthesis, microfluidic mixing and deposition

Additive manufacturing for biofabricated medical device applications

317

of scaffold materials, the use of small molecules or peptides to replace cost-prohibitive cell attachment, growth and differentiation factors, and related approaches. These will all need to be generated within an overarching quality system during product development. Use of advanced biomaterials may also necessitate the manufacturing of advanced 3D bioprinting systems to accommodate these materials and to enable scalability. Standardization of both the printer and the filaments present important opportunities for the field, leading to more rapid and reproducible results between laboratories and among manufacturing facilities and companies. Other opportunities for AM in the fields of medicine and biomedical science include moving from often wasteful, imprecise and poorly reproducible TM methods, to systems that are green, more sustainable, renewable, and of near-zero to zero-defect manufacturing. AM is a “green technology” in that nearly 100% of the material is used and little to no waste or byproducts are formed in the manufacturing processes, unlike classic subtractive (TM) methods. AM also provides a potent platform for generating patient specific products and enabling precision medicine, in place of common medical device approaches, where a few generic sized materials fit some patients (after nipping, tucking, or hammering, as is often the case).

1.5 Synopsis In this chapter, we describe the recent progress and current trends in using biological and synthetic materials for AM as may be applied for medical devices. Important topics in the field include the material synthesis and quality control needed for manufacturing mechanically strong filaments, fibers, or gels for biomedical applications and the printing of living scaffolds. Advanced biomedical AM applications in this chapter include the manufacturing of organoids, acellular scaffolds, cellularized devices, implantable and wearable medical devices and diagnostic tools, while also addressing manufacturing challenges for AM industrialization. Also, discussed in this chapter are issues in AM for medical devices relating to the needs for large scale manufacturing, high reproducibility, controlled cost of goods, and manufacturing of materials that will be used in a regulated market (e.g. by the FDA). Beyond the inorganic bioprinted materials, we discuss needs of living devices made by AM further have the omnipresent tissue engineering challenge of nutrient and waste exchange in “large” (>200 µm) constructs. Additionally, major challenges at the biological level that must be considered in AM including possible needs from storage and packaging of a viable graft with cells or of a protein-containing material, along with sterility and sterilization methods concerns, as also examined here.

2  Advanced biological materials in additive manufacturing 2.1 Introduction The expansion of available biomaterials—and tissue engineering techniques to generate scaffolds from the wide array of biomaterials—has largely driven the success of tissue engineering and regenerative medicine in preclinical studies, and, to a small

318

Additive Manufacturing

extent, for clinical products. Biomaterial scaffolds alone have shown marked therapeutic benefits, particularly when the materials integrate well into host tissue to allow for intrinsic patient cell infiltration. It may be possible to speed healing and better restore preinjury performance by using precision AM materials in place of bulk materials or TM produced methods, or precellularized scaffolds made via AM. Common material feedstocks used in AM include metals and plastics, biopolymers, cells and hydrogels (such as synthetic and natural extracellular matrices). AM feedstock materials may be prepared as powders, filaments, resins, melts or liquid solutions for assembly into intricate 2D or 3D materials. Many TM and AM processes require secondary manufacturing techniques. Techniques such as braiding, weaving, welding, coating, vapor deposition, and electroplating are used to build structures or to enhance the base structure’s characteristics, and are approaches that may also be applied to the 3D biomaterial AM, particularly for complex composite materials.

2.2  3D scaffold printing of biopolymers In 1986, Charles W. Hull described the 3D printer, based on a stereolithography (SLA) process wherein thin layers of material are cured by ultraviolet (UV) light to build 3D structures [6]. This fundamental concept of 3D printing has led to technologies for building 3D biological or living structure, commonly referred to as 3D bioprinting [2,7]. Rather than only depositing metals, polymers or other inorganic materials, a 3D bioprinter is designed to deposit living (typically eukaryotic) cells and/or a biocompatible matrix (e.g. a biopolymer) to provide structural support. 3D bioprinting typically involves the use of a “print head” in a deliberate X- and Y-axes coordinates, with cells, or “bioink” built layer-by-layer in the Z-axis via a movable stage or print bed. A timeline highlighting major events in 3D printing is depicted in .

2.2.1  Inkjet printing The 3D bioprinter print head classically uses common inkjet cartridges for cell and matrix deposition [8,9]. Inkjet, or drop-on-demand printers use acoustic (ultrasound) or thermal energy to deposit drops of water containing cells in 2D space via moving the cartridge fixed on a linear rail, commonly with a moving stage to add dimensionality to the prints. This so-called top-down technology was first used with off-the-shelf inkjet printers and cartridges commonly used for depositing ink on paper [10], with limited resolution for cell deposition in the 350 µm range in thick gels for patterning cells and biologically active proteins. However, resolution has improved greatly over the past decades and continues to improve. Recent advances include multilayer nanofilms developed by Choi et al. [11], which incorporated bFGF growth factor in a glycerol-based bioink to promote cell proliferation. Using microfluidics and piezoactuators to generate pulsed pressure in the cell delivery nozzle, Kim et al. [12] have advanced drop-on-demand inkjet-based printing of cells to the 30-µm, which is within the cell-level range of accuracy for many mammalian cell types. These technologies are both remarkable, in that they support high cell viability.

Additive manufacturing for biofabricated medical device applications

319

Inkjet bioprinting is proving great value for producing patterned arrays, such as lab-on-a-chip application for diagnostics and other platform technologies. Inkjet technology is still largely limited by the use of liquid printable materials, and is generally less advanced at forming vertical (3D) structures, and thus is not ideal for tissue engineering applications. Nonetheless, work is advancing to promote inkjet for the fabrication of relatively thin tissues, including the retina. Two of the 60 types of cells present in the retina, specifically retinal ganglion and glial cells of rat origin, were ordered using a piezoelectric-based print head, with minimal adverse events noted on cell health postprinting. Challenges noted in this approach are consistency in cell deposition, the complexity required for integrating the other cell types, for vascularizing the tissue construct, and for hardware issues relating to reliability related to cell sedimentation in the print head [13].

2.2.2 Extrusion Recent approaches for 3D bioprinting use microfluidic devices for microextrusion, such as via tubing or drawn glass needles [14] to precisely control the flow and layering of living cells. Extrusion is commonly driven by pneumatics [15–18], precision screws [14,19], worm drives, motor-driven spool or related engineering methods to eject material from a precision print head. Fused deposition modeling (FDM) is also commonly used for the extrusion of biomaterials to form fibers. FDM is often implemented via a heated nozzle which deposits a thermoplastic or other melted material from a filament into a defined space that cools and hardens to form a rigid 2D or 3D structure. FDM extrusion technology is found in common consumer-grade 3D printers (e.g. the MakerBot™ and Felix™ 3D printers) and has become popular in academic laboratories due to its relative ease of use, low cost and minimal complexity, making it a valuable basic research tool for student, though not professional, use. Many materials can be plotted or expelled as fibers by extrusion printers that are not readily printable with inkjet technologies, including hydrogels [20], biopolymers [21,22], cells and cell spheroids, bioactive glasses [23], and demineralized bone matrix (DBM) [24], among others. The deposition of thermoplastics (biopolymers) and biological particulates (e.g. tricalcium phosphate, bone matrix) commonly involves a heated nozzle to heat the support polymer past its crystal transition state to extrude the filament into the print-bead. The high heat used in FDM precludes concomitant cell deposition, so cells are typically added to the scaffold by a second extruder or by other means to prevent cell damage. Interestingly, processing FDM materials under nitrogen, in an oxygen-free environment, produces mechanically superior polymers [25], making this approach of high potential utility for scalable manufacturing where high reproducibility and consistency are of paramount importance.

2.3  Synthetic polymer printing AM research for tissue engineered medical device products (TEMPs) has most commonly involved either photopolymeric or thermoplastic biodegradable polymers. In particular, synthetic poly(lactides) (PLA) and its many copolymers, chemical

320

Additive Manufacturing

modifications, and derivations thereof (such as poly-d,l-lactide [PDLLA] and polylactic-co-glycolide [PLGA]) have been explored extensively for AM of cartilage [5] and bone repair indications [26–29]. PCL has also been researched extensively for bioprinting [23,24] yet, the published work to-date on PCL, PLA, and other synthetic polymers relies on commodity-grade polymers, whereas higher purity medical-grade materials, synthesized using Good Manufacturing Practices (GMP) need additional research and development. GMP-grade materials are critically important for product development, and, where possible, should be used in mid-to-late-stage research efforts, as the quality of the polymer will significantly affect the stability and in-life performance of the material [30–32]. However, for biological applications where clinical translation is not the goal, such as in basic research, common, frugal polymers, even nonbiocompatible polymers such as acrylonitrile butadiene styrene (ABS) [5] have shown promise. ABS was shown to support cartilage and nucleus pulposus cell viability, growth, and ECM deposition comparable to more common PLA, and suggests promise for future variants of ABS, such as ABS M30i, to be biocompatible.

2.4  Biological polymer printing AM applications of biological materials, such as ECM or natural materials with innate biological properties, have thus far been used less frequently compared with synthetic polymers. Biological polymers are more complex, more expensive, and typically have less desirable and less tailorable mechanical properties relative to synthetic polymers. Many current applications of biological scaffolds use additional polymer materials for structural support, including biological matrices in small amounts to enhance the biological properties of the base polymer. Nonmammalian ECM sources are also commonly used. For example, alginic acid, or alginate, a biofilm produced by bacteria and a major component of the cell wall in brown algae, is commonly used as a base printing material, due to its nature of thrift, well-known material properties, and convenience. Mammalian ECM materials have been used in combination with a carrier through a coaxial microfluidic print head to extrude photocurable polyethylene glycol (PEG)-fibrin and alginate to print 3D scaffolds filled with axially aligned C2C12 “myoblasts” [33], with the alginate being used as a temporary support hydrogel, upon which the UV-cured PEG-fibrin matrix is formed. After 3 weeks of culture, these printed constructs showed a high level of myogenic differentiation and histological organization, including sarcomeric alignment and spontaneous contractility in vitro, and showed ordered myofibril formation in vivo when subcutaneously implanted for 28 days. Alginate was used as a carrier for collagen and gelatin for 3D cell printing of human corneal epithelial cells by temperaturecontrolled extrusion to form 30 mm × 30 mm× 0.8 mm (8 layer thick) constructs. In this work, sodium citrate was used to dissolve the alginate over time to allow the cells to deposit new ECM, leaving collagen to support the cells, yielding a scaffold, which exhibited faster growth and more mature markers of differentiation. In another common strategy to introduce biological materials into AM, Mozdzen et al. incorporated bovine collagen and shark fin chondroitin sulfate onto ABS by lyophilizing the protein onto the nonbiocompatible plastic [34]. This approach formed

Additive manufacturing for biofabricated medical device applications

321

fibers with variable elasticities and porosities to show a highly tailorable material that can tether various growth factors. Approaches using ABS, while economical and using a readily available and printable material, are however not well suited for clinical use due to low compatibility in vivo. Recently, nanofibrillated cellulose bioinks with alginate and hyaluronic acid have shown to be printable by extrusion along with induced pluripotent stem cell (iPSC)-derived chondrocytes to print porous 7 mm × 7 mm × 1.2 mm thick cartilage mimetic [4]. This approach produced advanced cartilage-like hierarchical structures, as shown by histology of the printed structure. These materials also proved unstable, likely due to degradation by metalloproteases. Matrigel™ has also been used with agarose, a polymer derived from seaweed, to support intestinal epithelial cells [35] and cardiac progenitor cells [36] in 3D constructs. While promising for basic research, possible diagnostics, or “lab-on-chip” applications, these methods that incorporate cells into mechanically unstable hydrogels or nonbiocompatible polymers, and using cell types which cannot be used clinically, limit the utility of these technologies for regenerative medicine purposes. Future advancements will be needed to replace current approaches with superior materials made of clinical grade (GMP) materials to suffice the regulated environments, in which these materials are envisioned to be used for. Fabricating the scaffolds using more robust methods that produce mechanically sound constructs, not fragile hydrogels or nonbiocompatible plastics, will further be critical from a biomaterial development standpoint. A final limitation of note is the manufacturability of parts (scaffolds, dressings, tissues, etc.) in bulk, and of the appropriate scale for the intended therapeutic use. Most 3D bioprinted polymers, synthetic and biologic, are a few cubic millimeters is size, and thus scaled-up production methods must be developed to accommodate for larger materials, both for the clinical needs and for economic manufacturing of these devices.

2.5  Cell printing While biomaterial scaffolds alone have shown marked therapeutic benefits [37–39], the application of therapeutic cells to the material may improve healing rates or overall therapeutic efficacy as compared to passive cellular ingrowth into an acellular biomaterial. Aside from decellularized native tissue and organ matrices [40,41], common current tissue engineering methods are not very well suited for cellularization, as the tissue engineered materials are often too dense (i.e. by electrospinning), too fragile (hydrogel or self-assembling amphiphiles), or have microarchitectures dissimilar to native tissue (i.e. sponges). While these common methods of 3D scaffold manufacturing may produce structures that support cells, the spatial orientation of the cells and matrix are foreign to the tissue being repaired, lending to nonoptimal interfacing with host tissue and less successful healing. Major current thrusts to produce living grafts with cells in AM/biofabrication are proceeding down paths of (1) single cell resolution and multicellular cell aggregates, or (2) spheroids [42–44] deposited within or upon hydrogels or biopolymer fibrous scaffolds. Spheroids may be formed by the hanging-drop, small wells in a nonadherent

322

Additive Manufacturing

plate, or related methods. Cell spheroids can contain dozens, hundreds or thousands of cells, and are generally dispensed in a microliter or less of fluid, such as by dropon-demand systems. For basic research or diagnostic purposes, spheroids provide a shotgun-approach to deposing cells in 2D or 3D, such as into a hydrogel or even scaffold free by (using wash-away agarose supports) [19,42,45]. This approach allows for more rapid printing of many cells and increasing the cell density of a manufactured construct, yet spatial precision of individual cells is limited. Conversely, extrusion, laser or inkjet printing of near single-cell resolution is also possible. Laser-printing of single cells has been shown highly effective, with near one hundred percent viability [46]. Drop-on-demand (inkjet-based) approaches have also shown near single cell resolution and better fluid flow control by using bioinks containing cells in a gel or with surfactants [47,48], and delivering from a stand-alone print head that may be used in conjunction with a second print-head for matrix deposition. Single cell printing has theoretical advantages for studying some aspects of cell biology, and may prove useful in diagnostic applications. Gains in precision in single cell printing are, however, traded-off with losses in overall print speed, and with challenges to scale-up in 3D, where bulk cell deposition with spheroids, or conventional cell-seeding methods lend more favorable to 3D scale-up. Cell printing methods most appropriate for a given biofabricated medical device will ultimately be based on intended product features and specifications, as well as other manufacturing constraints (e.g. time, cost, complexity, and reproducibility). For biofabricated medical device production—as in cell therapy field in general— the cell type selection, donor selection, cellular health, and immune response of the chosen cells are important considerations. Using autologous cells may be feasible for some indications and circumvent immune complications, but impractical for others where cells must be expanded. Thus, allograft cells that have been expanded and characterized are preferred, yet selection of an ideal donor, and ensuring the quality and long-term health of those heavily expanded cells are also problematic. Highly expanded cells in culture are prone to genomic and phenotypic alternations, possibly including transformation to cancerous cells types. Additional challenges with cellularized grafts include storage or preservation and ensuring viability of any construct beyond 100–300 µm. Within these constructs, if the cells that have migrated into the graft are not supplied with blood flow they will rapidly become necrotic. Finally, with concerns of both particulate material contaminants, bacterial, fungal, viral, and prions in any living graft, engineering needs for producing equipment which can be validated for sterilization and can be cleaned between runs is also a critical need. Moving from the nano- or microscale through the mesoscale to macroscale products is particularly challenging and is an area not yet well addressed by the AM medical device community, where viable cells or biological polymers are anticipated for use. The printing of cells (around 30 µm in diameter) on fibers (around 100 nm to a few hundred microns in diameter) in three-dimensions to replicate even a simple tissue structure containing billions of cells and cubic centimeters of volume is limited currently both in terms of hardware and software. As computer processing has grown, so too need 3D bioprinting technologies to accommodate aspirations of medical device manufacturing.

Additive manufacturing for biofabricated medical device applications

323

3  Advanced biomanufacturing modalities for medical devices 3.1 Electrospinning Electrospinning is a fiber production method that has been explored for over a hundred years, since at least 1900, when John Francis Cooley’s filed the first patent in this field [49]. Electrospinning is an electrohydrodynamical process wherein a polymer solution is pumped in a syringe or otherwise presented in a controlled manner into a large static electric field (around 5–100 kV with negligible amperage). When sufficiently high voltage is applied to the liquid, electrostatic repulsions overcome the surface tension of the polymer, droplets of the solution are stretched and a stream of liquid polymer erupts from the surface. This eruption (the Taylor cone) elongates and, if the molecular cohesion of the liquid is sufficiently high, a liquid jet is produced. This jet flies at a high rate and continues to elongate as it moves toward a grounded or oppositely charged collector; the solvent evaporates and the polymer dries during the transit between emitter and collector. As the jet dries the electrical properties shift from ohmic to convective current flow, with the charge moving to the fibers surface. Elongation and related thinning of the polymer fibers are further driven by bending instabilities from the electrostatic repulsions [50,51]. The magnitude of the fiber elongation and thinning can be controlled to a large extent by altering processing parameters to produce fibers of variable geometries. The spatial patterning of the electrospun fibers can further be controlled by varying the electrostatic field by placing an auxiliary electrode either through placement of auxiliary electrodes in the electrostatic field (targeting) [52–57], by modifying the field vectors, or by alternating the current to produce an electrodynamic field (steering) [58]. Alternatively, engineered collecting surfaces, such as a high-speed rotating drum or wire wheel [59], spinning disk [60], or other surfaces can be used to drive the spatial patterning of the produced fibers. Electrospinning of polymers from liquid solutions is a processing technique that has been used widely in the generation of 3D micro- and nanofibrous scaffolds for tissue engineering applications. This technology allows formation of fibers with diameters ranging from tens of nanometers to several microns. The small fibers produced by electrospinning provide a large surface-to-volume ratio and an interconnected pore structure with high permeability, both of which are desirable in a biological setting [61–78]. The electrospinning process is not classically considered an AM technique due to the dynamic and often uncontrolled nature of fiber deposition [79]. However, the combination of electrospinning and AM approaches has been utilized to achieve fiber collection and deposition in an ordered and controlled fashion to generate hybrid-manufactured biomaterials [79–83]. Major advances and events in the history of electrospinning have been highlighted in Fig. 9.2. AM uses CAD to produce complex 3D structures in a layer-by-layer process; therefore, providing precise control over scaffold architecture [79–83]. A combination of electrospinning and FDM has recently been utilized to fabricate hybrid grafts of fibrous scaffolds lined or wrapped with a layer of printed polymer. Centola et al. developed a heparin functionalized poly-l-lactide (PLLA)/poly-ε-caprolactone (PCL) scaffold

324

Additive Manufacturing

Figure 9.2  Major events during the history of 3D printing and electrospinning.

for vascular tissue engineering by combining electrospinning and FDM. PLLA/heparin fibrous scaffolds were fabricated in tubular shape by electrospinning and further wrapped, on the outer layer, with a single coil of PCL formed by FDM technique. This scaffold design allowed the generation of both a drug delivery system and a microenvironment able to support human mesenchymal stem cells morphology, viability, and induced endothelial differentiation. Additionally, the PCL external coiling improved the mechanical strength of the microfibrous scaffold. Altogether, the development of a suitable device with a potential use in tissue engineering of the vascular grafts was made possible by combining electrospinning and FDM techniques [81]. Another AM approach integrating with electrospinning is the area of rapid prototyping. Rapid prototyping can provide precise, robust, and versatile ways of manufacturing collector plates with controlled microgeometries for electrospinning. These collector plates can be modified based on the user needs and topography of the target tissue and allow the formation of hybrid electrospun scaffolds with patterned fibers [80]. Rogers et al. [80] demonstrated that the geometry of the collector plate can influence the collection patterns of the fibers as well as fiber diameters within a single design without the need to alter any of the electrospinning conditions. As the structure of the nanofibrous scaffolds influences cellular response, generation of patterned scaffolds that closely replicate the structure of the native ECM of the target tissue is of great interest in the development of in vitro models. Rogers et al. [80] demonstrated that the sinusoidal electrospun scaffolds mimicking the fibrous nature and the native 3D geometry of the dermal papillae, provided a desirable microenvironment for dermal stem cell niche and supported cell adhesion, viability and proliferation in vitro, which potentially can be used as a powerful tool in disease modeling.

Additive manufacturing for biofabricated medical device applications

325

Melt electrospinning writing is another recent manufacturing approach that utilizes high viscosity, low conductivity fluids and allows fabrication of scaffolds with controlled fiber deposition and porosity [82]. Other electrospinning methods such as near-field electrospinning [84,85], high precision deposition electrospinning [86], and scanning tip electrospinning [87] have been used as an AM approach in the production of well-ordered nanofiber patterns.

3.2 Fabrics Smart fabrics and interactive medical textiles are a relatively new area of research, with potential biomedical applications. Manufacturing of smart textiles requires a complex and innovative technological approach, which combines conventional textile manufacturing such as weaving, knitting, and embroidery with technologies such as coating, lithography, and ink-jet printing [88]. Based on their application, smart fabrics involve materials and structures that respond to electrical, mechanical, chemical, thermal, optical, or magnetic stimuli [88]. Applications of smart textiles in medicine vary from sutures in surgical procedures to wound dressings to bio-monitoring sensors and wearables for therapy and wellness [88–90]. Depending on the type of wound or surgery site, degradable (temporary) or nondegradable (permanent) surgical sutures are typically used to hold the damaged tissue together during healing. Surgeries such as hernia repair, tendon repair, rotator cuff repair, and orthopedic applications, require strong mechanical support over a long period of time. However, depending on the type and number of sutures and the exertion of large tensile forces suture failure or tears may occur [91–94]. To prevent damage to the wound site as well as gather useful information during the healing process, monitoring tensile forces on the suture is critical [95]. To achieve this DeRouin et al. have designed a wireless sensor made of a thin magnetoelastic strip of Metglas 2826 MB, an amorphous ferromagnetic alloy designed by Metglas, Inc. This sensor is attached to suture thread on both ends and the tensile forces on the suture can be directly transferred to the sensor [95]. The use of such sensors in medical applications will allow optimization of surgical techniques and monitoring of tissue regeneration and regained strength of the damaged tissue. Using wearables for therapy and wellness is another medical device application of smart textiles. One example of their application is in patients with diabetes. Patients with diabetics have a high chance of developing foot ulcers, which is usually loss of tissue below the malleoli [96]. Untreated ulcers can develop infection and pose the risk of amputations [97]. It has been shown that the mainstay therapy for these foot ulcers is to reduce the mechanical load and stress [98]. In a recent study, Raviglione et al. developed a wearable system that can be used to monitor the stress applied to the ulcerated area, provide real-time feedback to the patient, and deliver data to the medical provider [96,99]. This system was developed by using a stretchable, bendable, washable, and sewable textile pressure sensor patented [99]. This sensor acts as a variable resistor (piezo-resistive), where the resistance is inversely proportional to the pressure. It is sewn to an elastic band, which adheres to the site of the ulcer. The sensor system collects, amplifies and filters the raw data, and transmits these data wirelessly via Bluetooth® Smart to a connected device using the internal Bluetooth module [96].

326

Additive Manufacturing

Additionally, textile-based drug delivery systems have been investigated through the combination of textile and biotechnology with chemistry and pharmacy [100]. Traditionally, wound dressings composed of a piece of material such as cloth, gauze, film, gel, or foams. More recently, functionalized materials with various coatings have been investigated more extensively [101]. Coated smart textiles for wound dressing are engineered to exhibit flexibility, strength, and be air and moisture permeable [101]. As an example, hydrocolloid dressings typically contain nonwoven polyester fibers as their textile element and are further functionalized with several components such as polyurethane gels, protein, and polysaccharides. These hydrocolloid dressings interact with the wound exudates and form a protective surface over the wound, which separates from the dressing upon removal, preventing damage to the newly formed skin [101]. Fiber optic-based sensors (FOSs) have been used extensively in the fields of engineering and medicine. These sensors have good accuracy, good sensitivity, large bandwidth, are immune to electromagnetic interferences, allow the measurement of physical and chemical parameters, and can both sense and transmit signals. These features make FOSs an emerging solution for the monitoring of physiological parameters and more generally for applications in medicine [102]. The immunity of these sensors to electromagnetic interferences makes them a great candidate technology to be used during MRI procedures as monitoring of the respiratory function of sedated or anesthetized patients during MRI is critical [103]. During the last decade, the integration of fiber optic technology and smart textiles has been explored substantially [104,105]. Fiber Bragg grafting sensors (FBGs) are one of the most frequently employed technologies in the design of smart textile based on fiber optics due to their high sensitivity to strain. FBGs are constructed in a short segment of optical fiber and consist of a periodic perturbation of the refractive index along the fiber core length obtained by exposure of the core to an intense optical interference pattern. These sensors reflect a narrow range of wavelengths and transmit all others [102,103]. FBG sensors have been used in monitoring respiratory movements [106] and during the last few years they have been embedded in the smart textiles. The high sensitivity of FBG-embedded smart textiles to strain allows the detection of small strains making them highly beneficial in monitoring the very small thoracic movements that are too small to be detected by other sensors. The group involved in the European project OFSETH (optical fiber sensors embedded into technical textile for healthcare) designed a novel noninvasive monitoring system based on FBG sensors to monitor both abdominal and thoracic respiratory movements. This textile based monitoring system offered comfort in addition to reliable sensitivity readings of between 0.1% and 5% elongation [103].

3.3 Molding Molding has been in use for millennia and is one of the oldest methods of manufacturing, yet we are still finding new uses due to its ability to make complicatedly shaped parts quickly and repeatedly. AM has provided new, impactful manners to create advanced molds such as: those used for the base material to create trabecular metal;

Additive manufacturing for biofabricated medical device applications

327

molds based on renderings from medical imaging scans; and dies with shapes not readily or economically manufactured by other methods. Trabecular metals, a successful technology based on molding, have advanced greatly over the last 20 years and have recently shown potential in manufacturing implantable medical devices. This material has shown potential for bone ingrowth into an engineered implant, reducing the need for additional fixation devices while creating a more stable bond between the implant and the host tissue [107]. Trabecular metal is made by compressing an open cell carbon sponge into a mold, followed by vapor depositing tantalum onto the material, and shaping to its finished dimensions with electrical discharge machining (EDM) [107]. This material has shown potential for bone ingrowth into an engineered implant, reducing the need for additional fixation devices while creating a more stable bond between the implant and the host tissue [108,109]. The open cell structure of the sponge provides pathways for ingrowth of the host tissue, while the tantalum provides the mechanical strength and chemical resistance to enable the parts to function in the body. EDM has shown to be the best method for machining the final shape because it is a noncontact method that uses electrical current to remove material, and tolerances of 0.01 mm are readily achieved; traditional machining methods such as milling and drilling tend to smear the material leading to occlusion of the pores [108]. In a novel hybrid-manufacturing approach (TM and AM), Rodriguez et al. [110] found that DBM fibers could be compression molded to create a mechanically stable and bioactive implant as a medical device for bone regeneration. In this work, 3D printed and machined molds were used to make standard-sized parts for spinal fusion surgery, wherein the DBM allografts (CNC cut bone fibers with surface patterning), were compressed in the molds to form stable, osteoinductive implants. They also showed 3D models from both patient CT images and 3D scanned donor tissues could be rendered via CAD/CAM from a patient’s specific defect to create 3D-printed molds. This process formed 100% biological parts, without filler or support material, precisely matching the patient’s anatomy or defect, such as producing custom parts for a cleft palate or acetabular cup hip repair implant. This technology has the potential to decrease surgical time and improve healing by providing surgeons with custom solutions to surgeries where loose, particulate granules of DBM or generic, nonosteoinductive filler materials (such as collagen sponges or tricalcium phosphate implants) would typically be used clinically. Nitinol stents are most commonly laser cut [111] from tubes made of nitinol made in vacuum induction melting. However, Mirizzi [112] developed an AM approach with a rotational molding method to create stress-free stents which are obtained directly from the base material. This advance manufacturing technology has the possibility of reducing production time and cost by eliminating the processes of manufacturing then cutting the tube to a finished shape. Most recently, a potential breakthrough has been made to 3D print a large diameter thermoplastic stent [113] that expands and contracts similarly to nitinol stents, allowing for minimally invasive implantation. The stent is manufactured of a polymer that is susceptible to hydrolytic breakdown allows for growth of the pediatric patient and may avoid the potential pitfall of restenosis found in persistent nitinol stents [114–116].

328

Additive Manufacturing

3.4 Microfluidics Microfluidics is defined as the precise spatial and temporal control of fluids and fluid reactions in systems that are on the submillimeter, or micro scale. A main benefit of microfluidic systems is the minimal use of reagents and streamlined flow paths caused by laminar flow, which is ideal for controlling sensitive chemical reactions, and manipulating cells. According to research by Warkiani et al. [117] microfluidic systems operated in parallel could replace membrane-based filters for microfiltration using hydrodynamic forces in a technique called inertial sorting [118,119]. Their proposed system that filters CHO and yeast cells, common cells types used for large-scale bioreactors, “can replace existing filtration membrane and provide passive (no external force fields), continuous filtration, thus eliminating the need for membrane replacement” and thus lends itself to high-throughput, low-cost manufacturing. Dielectrophoresis is a technique commonly used in microfluidic devices with embedded electrical contacts to sort cells based on their type and polarizability and provides similar manufacturing benefits to inertial sorting [120]. Rapid mechanical deformation of cells has emerged as a promising, vector-free method for intracellular delivery of macromolecules and nanomaterials [121]. Rapid deformation, or “cell squeezing” significantly improves cell colony formation compared to other methods such as electroporation and cell-penetrating peptide use thus increasing throughput and efficiency. By sheathing a core flow by combining separate flow paths, polymers can be formed and extruded depending on relative fluid velocities and viscosities in a method called hydrodynamic focusing [122–124]. Extrusion through microfluidic systems at lower temperatures enables cell deposition when combined with biopolymers such as gelatin, gelatin/chitosan, gelatin/alginate and gelatin/fibrinogen [125]. Droplet formation and manipulation via sheath flow and sorting techniques have the potential to automate many tedious and error-prone assays, with impact in medical diagnostics [126]. The low diffusion distances for chemical and heat transfer make physical processes occur more rapidly on the microscale thus reducing required volume and therefore cost. Chemical gradients are important for cell signaling, and microfluidic systems are used to provide the ideal environment for well-controlled studies. Compared to traditional platforms that generate gradients on the millimeter to centimeter scale, microfluidics can produce intercellular signaling gradients, such as cytokines, in the range of 1–100 µm, as is seen in vivo, along with producing controlled hydrodynamic and mass transport conditions [127]. These techniques and others [128] have significantly contributed to biological research and show that microfluidic-based manufacturing systems could reshape the medical field.

3.5  Pop-Up Book microelectromechanical systems Microelectromechanical systems (MEMS) are micro-devices with a variety of electrical and mechanical functions and are powerful tools for enabling the miniaturization of sensors, actuators, and systems [129–132]. So-called Pop-Up Book MEMS is a novel micro manufacturing technique that enables the fabrication of complex,

Additive manufacturing for biofabricated medical device applications

329

multifunctional electromechanical devices and is suitable for small medical and microsurgical applications [133,134]. This technique allows the creation of 3D multifunctional structures with various layers and joints that remain flat with limited out-of-plane complexity and can ‘pop-up’ from a 2D folded configuration [133]. Gafford et al. designed a force-sensing surgical grasper through Pop-Up Book MEMS manufacturing technique. This grasper was shown to detect distal loads and could differentiate between different loading conditions. This technology has the potential in minimally invasive surgical applications and provides physicians with more information that will enable them to perform safer and more effective procedures [133].

3.6  Soft robots 3D printing methods such as stereolithography (SLA), FDM and selective laser sintering are AM processes in which successive layers of material are deposited until a 3D solid object is created [130,135,136]. The critical feature of these rapid prototyping methods is the ability to create entire structures in one fabrication step [130]. SLA and related printing technologies have enabled the manufacturing of inorganic devices, such as microsurgical tools that can be used minimally invasively and soft robots with elastic properties akin to biological tissue, of great potential utility for medicine. Conventionally, robots have been made of stiff materials such as hard metals with rigid links and joints. Although these robots can be incredibly powerful and precise, they lack compliance and adaptability [137–140]. Therefore, generation of soft robots has been of great interest due their flexibility, robustness, and safety in human interactions [138–140]. Soft robots are mainly composed of intrinsically compliant and deformable materials such as silicon rubbers and can match the elasticity of the biological tissues [137–140]. The primary interaction of medical/soft robots is with soft materials such as skin, muscle, and internal organs. It is of great importance for these robots to possess similar compliance and mechanical rigidity of the natural tissue for even load distribution and minimal interfacial stress concentration [137]. More recently, 3D printing has been used as an advanced manufacturing method in the creation of controllable soft robots [139,140]. This method makes high-throughput prototyping possible by allowing rapid design iterations with no additional cost for increased morphological complexity [139]. Bartlett et al. have designed a combustionpowered jumping robot with a rigid core and soft exterior using a multimaterial 3D printer [139]. This robot is composed of two nested hemispheroids. The top hemispheroid is composed of nine different layers, which through a stepwise gradient, allow a modulus of elasticity that ranges over three orders of magnitude. This setup creates a structure that can transform from a highly flexible (rubber-like) state to a fully rigid (thermoplasticlike) state [139]. Using a similar method, Umedachi et al. designed a highly deformable soft robot inspired by the physical and locomotion properties of caterpillars [140]. The soft body of this robot allowed waves of deformation and produced complex and robust gaits such inching and crawling. A variable friction contact point was designed to allow the robot to switch friction of its legs with the ground according to its posture and shape resulting in locomotion [140]. These technologies may prove highly useful in medical tool generation, particularly in minimally invasive surgery applications.

330

Additive Manufacturing

3.7 Metals AM of metals for medical applications has increased in popularity in the last decade mainly due to an increase in the number of printable materials and a decrease in the cost of commercial direct metal laser sintering (DMLS). This manufacturing method is simple; metal powder is blown into the path of a high-power laser which forms liquid beads and adheres to the surface of the target material. The materials that are commercially available now that are of high interest to the medical field are stainless steel, titanium and cobalt chrome alloys. These materials are high strength, commercially affordable and do not typically illicit an inflammatory response when implanted (depending on grade and on surface treatment). DMLS parts have incredible mechanical strength compared to other items produced using AM approaches, and these parts can be tailored to make implants based on scanned anatomical features for traumatic injury repair. DMLS metal parts have become clinically popular in the custom implants market, particularly with dental implants [141], spine fusion [142], facial reconstruction [143,144], and joint repair [145]. The metals can be porous when printed and help encourage bone ingrowth, which increases the strength of the bond [108,146,147]. Porosity also decreases the mechanical strength of the implanted material and when controlled, these materials can match the biomechanical properties of bone which helps prevent stress shielding and bone reabsorption after implantation.

4  Production challenges 4.1 Introduction Transitioning from parts made in a laboratory to the production of medical devices intended for distribution in a regulated (FDA) market presents myriad challenges beyond those of basic bench and pilot-scale research. Challenges include—yet are by no means not limited to—defining the methods to scale up or scale out the process, validating the process and the performance of the manufactured product, meeting regulatory requirements, and ensuring that the packaged part is sterile and of consistent quality from part-to-part and lot-to-lot. In this section, current and foreseeable manufacturing challenges to AM of medical devices are introduced. A simplified workflow for moving an AM medical device through research to formal product development, to production and logistic of launch along with regulatory tasks in the process is illustrated below (Fig. 9.3).

4.2  Sterility and packaging Traditional AM approaches, such as fused layer deposition of thermoplastic feedstock, produce inherently sterile parts due to the high temperature required to melt the material [148]. Recently, advances in printing soft tissue constructs have received attention because of their theoretical potential to significantly reduce cost and waiting time for tissue and organ recipients [2]; however, cells do not survive prolonged exposure to

Additive manufacturing for biofabricated medical device applications

331

Figure 9.3  Major events during medical device manufacturing.

high temperatures, so alternative 3D printing technology has been developed to form and fuse a biocompatible cellular matrix (bioprinting), such as contact/contactless bioprinting and direct photopolymerization [149–153]. In the AM of medical devices for human therapeutic indications without cells, special precautions still must be taken to prevent microbial contamination. The FDA has guidance on the sterilization or sterile processing approaches that are recommended, depending on the product or application [154]. Wherever possible, equipment should be made of materials designed for sterilization or disinfection, such as stainless steel, polypropylene, Teflon™ or other materials that can be sprayed or wiped with cleaning agents like bleach or alcohol before use. Both the manufacturing equipment and, where possible, the manufactured part surfaces should be devoid of crevices, cracks or holes with tortuous paths to help eliminate contaminants while cleaning. Furthermore, the number of moving parts in the device must be minimized to avoid disturbing airflow and creating debris, which is particularly challenging in many AM approaches. In 3D bioprinting with living cells, the cells, media, and syringes should all be sterilized and loaded into the device in as clean an environment as possible to avoid contamination. Work with live cells preferably is performed in a Class II or better biosafety cabinet or a clean room, and if possible in an entirely sealed system with extensive monitoring of the culture system in place. The FDA has published a technical document on these and other considerations for AM devices [155].

332

Additive Manufacturing

An often overlooked but equally important component of manufacturing after the product has been made is packaging. In a practical sense, the packaging is the first thing the surgical team will encounter, so it is important that extensive testing is done to ensure it opens easily and the product can be presented to the sterile field without difficulties [156]. In addition, the packaging must maintain the sterility of the product over an extended shelf life (from hours to days and up to several years in some cases), and allow for shipping in cold and warm conditions, and sustain integrity of the packaging and of the AM produced medical device inside through bumpy delivery truck rides and plane cargo bays. For 3D printed living tissues, packaging will need to sustain living or cryopreserved cells, which will have very specific needs for the packaging material, such as a thin wall for favorable heat transfer properties for cryopreservation. The shipping container, if needed, should be validated to ensure temperatures are maintained under worst-case conditions. The considerations for AM parts are ultimately no different than other medical devices. If the product is sterile after manufacturing, the packaging can be presterilized and sealed with the product in a sterile environment. If the product needs to be sterilized postpackaging, then the packaging will need to be able to survive the sterilization method. Several commercial packaging materials, for example Tyvek™, are available that allow steam or ethylene oxide to penetrate, but not microorganisms. Regarding sterilization methods, while markedly few methods have been proposed for sterilizing or disinfecting viable tissue (such as cold plasma), many methods are suggested by the Center for Disease Control (CDC) that can be applied for nonviable medical devices [157]. However, many medical devices that may be made by AM methods produce materials susceptible to damage through chemical or physical alteration from the sterilization process [158]. For example, while being excellent at obtaining a high Sterility Assurance Level (SAL), gamma irradiation has been shown to modify PCL to improve hydrophilicity at a low dose, yet high doses weaken the material properties [159] and can melt polymers at the nano- and micro-levels, significantly effecting the topology of the material. Electron beam (e-beam) sterilization typically uses less energy compared to gamma irradiation, yet e-beam may also adversely change the material properties of any AM product. It was recently shown that PLGA nanofibers reduced molecular weight with increasing e-beam dose, and the mechanical properties of the fibers were decreased in step [160]. Cell attachment and proliferation on e-beam treated matter were not apparently affected, however. Chemical methods, such as ethylene oxide, have been used widely in medical device manufacturing as well, yet such methods may also adversely alter the properties of the AM device, even for biological materials [161] as well as polymers. Thus, due care must be taken in choosing the appropriate sterilization methods, with the AM medical device in the final packaging, and testing of the performance specifications must further be validated post sterilization for any product that is terminally sterilized. Manufacturing of a commercial medical device does not necessarily require terminal sterilization, and in some cases, a device may be sterile processed or otherwise validated to be of low risk for microbial contaminations in the final packaged product. This decision will be driven by data, acceptable risk for the company, and by in many cases by the final ruling of the FDA.

Additive manufacturing for biofabricated medical device applications

333

4.3  Manufacturing challenges Plunging past the depths of research investigation to an interrogation of the process and specifics of the idealized product, product development presents a unique set of learning and challenges. Product development is a distinct move from research. In research, the manufacturing of a material, scaffold, or medical device is carried out in limited quantities, in a constantly evolving, iterative learning process, with work performed by highly skilled scientists or engineers. Development entails a transition to a robust and reproducible mode of manufacturing, producing a process which is like, but rarely identical to, the research method. In development, the scale of manufacturing grows by many folds, and the specifications, tolerances, or qualities of the product and the process are tightly controlled. Development requires determining the methods and equipment that will be used to manufacture suitable quantities of material, instituting quality control measures, such as setting product lot release criteria, and following the regulatory requirements of the region where the product is targeted for marketing. Development of an AM medical device cannot truly begin until the design is frozen. This includes determining the final shape, composition, and functional features, or product specifications of the device. Product specifications and actual prototypes of the mature research product should be reviewed by medical professionals, such as surgeons, to provide critical feedback on the envisioned product to help guide the development process. Consulting with end users before the completion of research will help to ensure the product has a niche in the market—small or large—and meets their specific needs, which are often not captured in any publication or from any expert lecture. In a direct conversation with the end user (surgeons, nurses, etc.) it is important to probe for in depth reviews of the product and gather input from multiple sources, as opinions frequently conflict about what the final design should be, and truly what are the real needs in the field . With key opinion leader feedback on the envisioned medical device gathered, guidance documents from the FDA (21 CFR 820.30) outline the design inputs and outputs for developing such a device, as well as how fulfilling the inputs will be measured through verification and validation (commonly referred to as a DIOVV). A single table with columns for user needs, design inputs, design outputs, verifications, and validations is typically created by industrial product development teams. This table will guide the process to ensure management, and all members of the team have the same goal in mind, and that the final product aligns with the design intent. A related form that may be created for guidance is a Target Product Profile. Risk management should also be conducted at this time. A Failure Modes Effects and Analysis form is a productive way of organizing the risks the product can have on patients, users, and staff; where those risks come from; and what can be done to mitigate them. In many cases for defining a successful product development, and for an FDA approval, animal studies may be required. A pilot animal study with enough animals to gather valuable information is conducted to select leading variations of the product as the final stage of research. Pilot studies should still adhere to strict statistical guidelines, with the number of animals to be tested, in so far as possible, determined by a power analysis. The standard research pilot testing run of an arbitrary or limited number of animals is likely to prove fruitless because perhaps there is a high probability that an

334

Additive Manufacturing

animal may die, or sample implant or recovery can be mishandled, and the results will not be strongly supported as statistically significant. It is not reasonable to discount a failing sample as an outlier, when there are only two implants that yielded good results. A representative device will need to be created for ISO 10993 and animal studies that are used to prove the device’s safety and efficacy to the FDA. This includes the product’s final formulation, packaging, and sterilization. The FDA encourages meeting with them to discuss your device and the methods you plan to use to test it. However, the meeting is only for specific questions, not general FDA procedures. The meeting also requires that submission of questions at the time the meeting is requested, which is 90 days in advance, so this time must be accounted for this in the product development timeline. In an AM, as in a TM process, a robust manufacturing system is required to meet the device’s design inputs, which must include the tolerances of the parts. Manufacturing engineers can work with the design inputs to define the process and specify the equipment necessary to scale up and manufacture the device. All requirements of the device and equipment need to be fully understood before purchasing or building custom machinery. For example, special precaution will need to be taken if the ultrafine particles created during 3D printing [162] will have an impact on the cleanroom environment as per ISO 14644-1 standards. Process validations of AM produced devices following the guidelines of FDA 21 CFR 820.75 must be conducted to ensure the manufacturing equipment can repeatedly make the devices in a stable and capable process. In short, the equipment must be installed and confirmed to continually be making parts within acceptable specifications. AM techniques, such as electrospinning and 3D printing, have characteristics that are not as well controlled and common as the more TM technologies. However, if the machinery is properly maintained, the staff is properly trained, and the expectations are within the bounds of the technologies’ intended use, these advanced AM processes should be able to conform to the same process validation guidelines and have market potential to contribute to medical device manufacturing [163–169]. Finally, after the device is proven to be safe and effective, the process is shown to be in control, and risks have been documented and mitigated, a submission can be made to the FDA. The review process typically takes 90 calendar days to complete. A request may be made for additional information or data for clearance of the medical device, or a device may be cleared for market after this review.

5 Summary AM methodologies have vast potential for advancing tissue engineering, at last, into medical device products that are useful to people. AM approaches are now being used to build scaffolds, spacers, sensors, and structural materials for medical applications, and more. AM has great potential for medical device companies in the immediate future, particularly for manufacturing devices for the spine, orthopedic, and dental indications. Great challenges remain for the production of viable grafts from structural proteins (collagen and other ECMs) by AM to produce functional tissues and organs as medical devices, yet the respective and extraordinary interests in the field from researchers, the biotech industry sector, and the public are providing tremendous thrust for advancing this field.

Additive manufacturing for biofabricated medical device applications

List of acronyms 2D 3D ABS AM ASTM CAD CAE CAM CDC CT DBM DHF DIOVV DMLS ECM EDM FBG FDA FDM FOS GMP iPSC ISO µm MEMs mg ml mm MPa MRI nm PCL PDLLA PDMS PEG PLA PLGA PMA PTFE QC SAL SEM SLA SOP TEMP TM

two-dimensional three-dimensional acrylonitrile butadiene styrene additive manufacturing American Section of the International Association for Testing Materials computer-assisted design computer-assisted engineering computer-assisted modeling center for disease control computed tomography demineralized bone matrix design history files design inputs, outputs, verifications, and validations direct metal laser sintering extracellular matrix electrical discharge machining fiber Bragg grafting Food and Drug Administration fused deposition modeling fiber optic-based sensors good manufacturing practices induced pluripotent stem cell International Standards Organization micrometer microelectromechanical systems milligram milliliter millimeter megapascal magnetic resonance Imaging nanometer polycaprolactone poly-d,l-lactide polydimethylsiloxane polyethylene glycol polylactic acid polylactic-co-glycolide premarket approval polytetrafluoroethylene quality control sterility assurance level scanning electron microscopy stereolithography standard operating procedures tissue engineered medical product traditional manufacturing

335

336

Additive Manufacturing

References [1] Itoh M, Nakayama K, Noguchi R, Kamohara K, Furukawa K, Uchihashi K, et al. Scaffold-free tubular tissues created by a bio-3D printer undergo remodeling and endothelialization when implanted in rat aortae. PLOS ONE 2015;10:e0136681. [2] Murphy SV, Atala A. 3D bioprinting of tissues and organs. Nat. Biotechnol. 2014;32: 773–85. [3] Bianconi E, Piovesan A, Facchin F, Beraudi A, Casadei R, Frabetti F, et al. An estimation of the number of cells in the human body. Ann. Hum. Biol. 2013;40:463–71. [4] Nguyen D, Hägg DA, Forsman A, Ekholm J, Nimkingratana P, Brantsing C, et al. Cartilage tissue engineering by the 3D bioprinting of iPS cells in a nanocellulose/alginate bioink. Sci. Rep. 2017;7:658. [5] Rosenzweig DH, Carelli E, Steffen T, Jarzem P, Haglund L. 3D-printed ABS and PLA scaffolds for cartilage and nucleus pulposus tissue regeneration. Int. J. Mol. Sci. 2015;16:15118–35. [6] C.W. Hull. U.S. Patent No. 4,575,330, 11 March, 1986. [7] Nakamura M, Iwanaga S, Henmi C, Arai K, Nishiyama Y. Biomatrices and biomaterials for future developments of bioprinting and biofabrication. Biofabrication 2010;2:014110. [8] Cui X, Boland T, D D’Lima D, K Lotz M. Thermal inkjet printing in tissue engineering and regenerative medicine. Recent Pat. Drug Deliv. Formul. 2012;6:149–55. [9] Xu T, Kincaid H, Atala A, Yoo JJ. High-throughput production of single-cell microparticles using an inkjet printing technology. J. Manuf. Sci. Eng. 2008;130:021017. [10] Roth EA, Xu T, Das M, Gregory C, Hickman JJ, Boland T. Inkjet printing for highthroughput cell patterning. Biomaterials 2004;25:3707–15. [11] Choi M, Park HH, Choi D, Han U, Park TH, Lee H, et al. Multilayer nanofilms via inkjet printing for stabilizing growth factor and designing desired cell developments. Adv. Healthcare Mater. 2017;6. [12] Kim YK, Park JA, Yoon WH, Kim J, Jung S. Drop-on-demand inkjet-based cell printing with 30-µm nozzle diameter for cell-level accuracy. Biomicrofluidics 2016;10:064110. [13] Lorber B, Hsiao W, Hutchings IM, Martin KR. Adult rat retinal ganglion cells and glia can be printed by piezoelectric inkjet printing. Biofabrication 2013;6:015001. [14] Reid JA, Mollica PA, Johnson GD, Ogle RC, Bruno RD, Sachs PC. Accessible bioprinting: adaptation of a low-cost 3D-printer for precise cell placement and stem cell differentiation. Biofabrication 2016;8:025017. [15] Khalil S, Sun W. Biopolymer deposition for freeform fabrication of hydrogel tissue constructs. Mater. Sci. Eng. C 2007;27:469–78. [16] Chang CC, Boland ED, Williams SK, Hoying JB. Direct write bioprinting three-dimensional biohybrid systems for future regenerative therapies. J. Biomed. Mater. Res. B: Appl. Biomater. 2011;98:160–70. [17] Fedorovich NE, Swennen I, Girones J, Moroni L, Van Blitterswijk CA, Schacht E, et al. Evaluation of photocrosslinked lutrol hydrogel for tissue printing applications. Biomacromolecules 2009;10:1689–96. [18] Chang R, Nam J, Sun W. Effects of dispensing pressure and nozzle diameter on cell survival from solid freeform fabrication-based direct cell writing. Tissue Eng. A 2008;14: 41–8. [19] Jakab K, Damon B, Neagu A, Kachurin A, Forgacs G. Three-dimensional tissue constructs built by bioprinting. Biorheology 2006;43:509–13.

Additive manufacturing for biofabricated medical device applications

337

[20] Zhao Y, Li Y, Mao S, Sun W, Yao R. The influence of printing parameters on cell survival rate and printability in microextrusion-based 3D cell printing technology. Biofabrication 2015;7:045002. [21] Lei M, Wang X. Biodegradable polymers and stem cells for bioprinting. Molecules 2016;21:539. [22] Poh PS, Chhaya MP, Wunner FM, De-Juan-Pardo EM, Schilling AF, Schantz J, et al. Polylactides in additive biomanufacturing. Adv. Drug Deliv. Rev. 2016;107:228–46. [23] Nyberg E, Rindone A, Dorafshar A, Grayson WL. Comparison of 3D-printed poly-εcaprolactone scaffolds functionalized with tricalcium phosphate, hydroxyapatite, biooss, or decellularized bone matrix. Tissue Eng. A 2017;23:503–14. [24] Rindone AN, Nyberg E, Grayson WL. 3D-printing composite polycaprolactone-decellularized bone matrix scaffolds for bone tissue engineering applications. Methods Mol. Biol. 2017;. [25] Lederle F, Meyer F, Brunotte G, Kaldun C, Hübner EG. Improved mechanical properties of 3D-printed parts by fused deposition modeling processed under the exclusion of oxygen. Prog. Addit. Manuf. 2016;1:3–7. [26] Bose S, Vahabzadeh S, Bandyopadhyay A. Bone tissue engineering using 3D printing. Mater. Today 2013;16:496–504. [27] Ge Z, Tian X, Heng BC, Fan V, Yeo JF, Cao T. Histological evaluation of osteogenesis of 3D-printed poly-lactic-co-glycolic acid (PLGA) scaffolds in a rabbit model. Biomed. Mater. 2009;4:021001. [28] Kim H, Lee H, Chun G. Bioactivity and osteoblast responses of novel biomedical nanocomposites of bioactive glass nanofiber filled poly (lactic acid). J. Biomed. Mater. Res. A 2008;85:651–63. [29] Noh K, Lee H, Shin U, Kim H. Composite nanofiber of bioactive glass nanofiller incorporated poly (lactic acid) for bone regeneration. Mater. Lett. 2010;64:802–5. [30] Cam D, Marucci M. Influence of residual monomers and metals on poly (l-lactide) thermal stability. Polymer 1997;38:1879–84. [31] Södergård A, Näsman JH. Melt stability study of various types of poly (l-lactide). Ind. Eng. Chem. Res. 1996;35:732–5. [32] Wachsen O, Platkowski K, Reichert K. Thermal degradation of poly-l-lactide—studies on kinetics, modelling and melt stabilisation. Polym. Degrad. Stab. 1997;57:87–94. [33] Costantini M, Testa S, Mozetic P, Barbetta A, Fuoco C, Fornetti E, et al. Microfluidicenhanced 3D bioprinting of aligned myoblast-laden hydrogels leads to functionally organized myofibers in vitro and in vivo. Biomaterials 2017;131:98–110. [34] Mozdzen LC, Rodgers R, Banks JM, Bailey RC, Harley BA. Increasing the strength and bioactivity of collagen scaffolds using customizable arrays of 3D-printed polymer fibers. Acta Biomater. 2016;33:25–33. [35] Fan R, Piou M, Darling E, Cormier D, Sun J, Wan J. Bio-printing cell-laden Matrigel– agarose constructs. J. Biomater. Appl. 2016;31:684–92. [36] Gaetani R, Doevendans PA, Metz CH, Alblas J, Messina E, Giacomello A, et al. Cardiac tissue engineering using tissue printing technology and human cardiac progenitor cells. Biomaterials 2012;33:1782–90. [37] Capito AE, Tholpady SS, Agrawal H, Drake DB, Katz AJ. Evaluation of host tissue integration, revascularization, and cellular infiltration within various dermal substrates. Ann. Plast. Surg. 2012;68:495–500. [38] Hughes OB, Rakosi A, Macquhae F, Herskovitz I, Fox JD, Kirsner RS. A review of cellular and acellular matrix products: indications, techniques, and outcomes. Plast. Reconstr. Surg. 2016;138:138S–47S.

338

Additive Manufacturing

[39] Pittman TA, Fan KL, Knapp A, Frantz S, Spear SL. Comparison of different acellular dermal matrices in breast reconstruction: the 50/50 study. Plast. Reconstr. Surg. 2017;139:521–8. [40] Bulysheva AA, Burcus N, Lundberg C, Edelblute CM, Francis MP, Heller R. Recellularized human dermis for testing gene electrotransfer ex vivo. Biomed. Mater. 2016;11:035002. [41] Ott HC, Matthiesen TS, Goh S, Black LD, Kren SM, Netoff TI, et al. Perfusion-decellularized matrix: using nature’s platform to engineer a bioartificial heart. Nat. Med. 2008;14:213. [42] Gutzweiler L, Kartmann S, Troendle K, Benning L, Finkenzeller G, Zengerle R, et al. Large scale production and controlled deposition of single HUVEC spheroids for bioprinting applications. Biofabrication 2017;9:025027. [43] Mehesz AN, Brown J, Hajdu Z, Beaver W, da Silva J, Visconti R, et al. Scalable robotic biofabrication of tissue spheroids. Biofabrication 2011;3:025002. [44] Mironov V, Visconti RP, Kasyanov V, Forgacs G, Drake CJ, Markwald RR. Organ printing: tissue spheroids as building blocks. Biomaterials 2009;30:2164–74. [45] Norotte C, Marga FS, Niklason LE, Forgacs G. Scaffold-free vascular tissue engineering using bioprinting. Biomaterials 2009;30:5910–7. [46] Ringeisen BR, Kim H, Barron JA, Krizman DB, Chrisey DB, Jackman S, et al. Laser printing of pluripotent embryonal carcinoma cells. Tissue Eng. 2004;10:483–91. [47] Ferris CJ, Gilmore KJ, Beirne S, McCallum D, Wallace GG. Bio-ink for on-demand printing of living cells. Biomater. Sci. 2013;1:224–30. [48] Xu T, Jin J, Gregory C, Hickman JJ, Boland T. Inkjet printing of viable mammalian cells. Biomaterials 2005;26:93–9. [49] J. Cooley, Improved methods of and apparatus for electrically separating the relatively volatile liquid component from the component of relatively fixed substances of composite fluids. United Kingdom Patent 1900;6385:19. [50] Doshi J, Reneker DH. Electrospinning process and applications of electrospun fibers. J. Electrostat. 1995;35:151–60. [51] Li D, Xia Y. Electrospinning of nanofibers: reinventing the wheel? Adv. Mater. 2004;16:1151–70. [52] Alternative Methods for Fabricating Aligned Electrospun Fibers. http://electrospintech. com/alt_align_fiber.html#referencelist. [53] Carnell LS, Siochi EJ, Holloway NM, Stephens RM, Rhim C, Niklason LE, et al. Aligned mats from electrospun single fibers. Macromolecules 2008;41:5345–9. [54] Kiselev P, Rosell-Llompart J. Highly aligned electrospun nanofibers by elimination of the whipping motion. J. Appl. Polym. Sci. 2012;125:2433–41. [55] Scott-Carnell LA, Stephens RM, Holloway NM, Rhim C, Niklason L, Clark RL, et al. Method and system for aligning fibers during electrospinning 2011. [56] Theron A, Zussman E, Yarin A. Electrostatic field-assisted alignment of electrospun nanofibres. Nanotechnology 2001;12:384. [57] Wincheski RA. Method for predicting and optimizing system parameters for electrospinning system 2011. [58] Grasl C, Arras MM, Stoiber M, Bergmeister H, Schima H. Electrodynamic control of the nanofiber alignment during electrospinning. Appl. Phys. Lett. 2013;102:053111. [59] Katta P, Alessandro M, Ramsier R, Chase G. Continuous electrospinning of aligned polymer nanofibers onto a wire drum collector. Nano Lett. 2004;4:2215–8. [60] Zussman E, Theron A, Yarin A. Formation of nanofiber crossbars in electrospinning. Appl. Phys. Lett. 2003;82:973–5.

Additive manufacturing for biofabricated medical device applications

339

[61] Alessandrino A, Marelli B, Arosio C, Fare S, Tanzi MC, Freddi G. Electrospun silk fibroin mats for tissue engineering. Eng. Life Sci. 2008;8:219–25. [62] Barnes CP, Sell SA, Boland ED, Simpson DG, Bowlin GL. Nanofiber technology: designing the next generation of tissue engineering scaffolds. Adv. Drug Deliv. Rev. 2007;59:1413–33. [63] Boland ED, Matthews JA, Pawlowski KJ, Simpson DG, Wnek GE, Bowlin GL. Electrospinning collagen and elastin: preliminary vascular tissue engineering. Front. Biosci. 2004;9:1422–32. [64] Boland ED, Coleman BD, Barnes CP, Simpson DG, Wnek GE, Bowlin GL. Electrospinning polydioxanone for biomedical applications. Acta Biomater. 2005;1: 115–23. [65] Boudriot U, Dersch R, Greiner A, Wendorff JH. Electrospinning approaches toward scaffold engineering—a brief overview. Artif. Organs 2006;30:785–92. [66] Chen J, Yan C, Zhu M, Yao Q, Shao C, Lu W, et al. Electrospun nanofibrous SF/P(LLA-CL) membrane: a potential substratum for endothelial keratoplasty. Int. J. Nanomed. 2015;10:3337–50. [67] Chen J, Chang G, Chen J. Electrospun collagen/chitosan nanofibrous membrane as wound dressing. Colloids Surf. Physicochem. Eng. Aspects 2008;313–314:183–8. [68] Chew SY, Wen Y, Dzenis Y, Leong KW. The role of electrospinning in the emerging field of nanomedicine. Curr. Pharm. Des. 2006;12:4751–70. [69] Chong EJ, Phan TT, Lim IJ, Zhang YZ, Bay BH, Ramakrishna S, et al. Evaluation of electrospun PCL/gelatin nanofibrous scaffold for wound healing and layered dermal reconstitution. Acta Biomater. 2007;3:321–30. [70] Ghasemi-Mobarakeh L, Prabhakaran MP, Morshed M, Nasr-Esfahani M, Ramakrishna S. Electrospun poly(ε-caprolactone)/gelatin nanofibrous scaffolds for nerve tissue engineering. Biomaterials 2009;29:4532–9. [71] Jin H, Chen J, Karageorgiou V, Altman GH, Kaplan DL. Human bone marrow stromal cell responses on electrospun silk fibroin mats. Biomaterials 2004;25: 1039–47. [72] Kanani AG, Bahrami SH. Review on electrospun nanofibers scaffold and biomedical applications. Trends Biomater. Artif. Organs 2010;24:93–115. [73] Khil M, Cha D, Kim H, Kim I, Bhattarai N. Electrospun nanofibrous polyurethane membrane as wound dressing. J. Biomed. Mater. Res. B: Appl. Biomater. 2003;67: 675–9. [74] Ki CS, Park SY, Kim HJ, Jung HM, Woo KM, Lee JW, et al. Development of 3- D nanofibrous fibroin scaffold with high porosity by electrospinning: implications for bone regeneration. Biotechnol. Lett. 2008;30:405–10. [75] Kumbar SG, Nukavarapu SP, James R, Nair LS, Laurencin CT. Electrospun poly(lactic acid-co-glycolic acid) scaffolds for skin tissue engineering. Biomaterials 2008;29: 4100–7. [76] Lannutti J, Reneker D, Ma T, Tomasko D, Farson D. Electrospinning for tissue engineering scaffolds. Mater. Sci. Eng. C 2007;27:504–9. [77] Matthews JA, Wnek GE, Simpson DG, Bowlin GL. Electrospinning of collagen nanofibers. Biomacromolecules 2002;3:232–8. [78] Matthews JA, Boland ED, Wnek GE, Simpson DG, Bowlin GL. Electrospinning of collagen type II: a feasibility study. J. Bioact. Compat. Polym. 2003;18:125–34. [79] Dalton PD, Vaquette C, Farrugia BL, Dargaville TR, Brown TD, Hutmacher DW. Electrospinning and additive manufacturing: converging technologies. Biomater. Sci. 2013;1:171–85.

340

Additive Manufacturing

[80] Rogers CM, Morris GE, Gould TW, Bail R, Toumpaniari S, Harrington H, et al. A novel technique for the production of electrospun scaffolds with tailored three-dimensional micro-patterns employing additive manufacturing. Biofabrication 2014;6:035003. [81] Centola M, Rainer A, Spadaccio C, De Porcellinis S, Genovese J, Trombetta M. Combining electrospinning and fused deposition modeling for the fabrication of a hybrid vascular graft. Biofabrication 2010;2:014102. [82] Hochleitner G, Jüngst T, Brown TD, Hahn K, Moseke C, Jakob F, et al. Additive manufacturing of scaffolds with sub-micron filaments via melt electrospinning writing. Biofabrication 2015;7:035002. [83] Quan Z, Wu A, Keefe M, Qin X, Yu J, Suhr J, et al. Additive manufacturing of multi-directional preforms for composites: opportunities and challenges. Mater. Today 2015;18: 503–12. [84] Sun D, Chang C, Li S, Lin L. Near-field electrospinning. Nano Lett. 2006;6:839–42. [85] Chang C, Limkrailassiri K, Lin L. Continuous near-field electrospinning for large area deposition of orderly nanofiber patterns. Appl. Phys. Lett. 2008;93:123111. [86] Hellmann C, Belardi J, Dersch R, Greiner A, Wendorff J, Bahnmueller S. High precision deposition electrospinning of nanofibers and nanofiber nonwovens. Polymer 2009;50:1197–205. [87] Kameoka J, Orth R, Yang Y, Czaplewski D, Mathers R, Coates GW, et al. A scanning tip electrospinning source for deposition of oriented nanofibres. Nanotechnology 2003;14:1124. [88] Mecnika V, Hoerr M, Krievins I, Schwarz A. Smart textiles for healthcare: applications and technologies. 2014. [89] Langenhove LV. Smart Textiles Medicine and healthcare. Mater. Syst. Appl. 2007; 106-122:153–300. [90] Coyle S, Diamond D, Kirstein T. Medical Applications of Smart Textiles. Multidisciplinary Know-How for Smart-Textiles Developers. Cambridge, UK: Woodhead Publishing Ltd.; 2013. 420-443. [91] Barber FA, Cawley P, Prudich JF. Suture anchor failure strength—an in vivo study. Arthroscopy 1993;9:647–52. [92] Bynum CK, Lee S, Mahar A, Tasto J, Pedowitz R. Failure mode of suture anchors as a function of insertion depth. Am. J. Sports Med. 2005;33:1030–4. [93] Cummins CA, Murrell GA. Mode of failure for rotator cuff repair with suture anchors identified at revision surgery. J. Shoulder Elbow Surg. 2003;12:128–33. [94] Meyer DC, Nyffeler RW, Fucentese SF, Gerber C. Failure of suture material at suture anchor eyelets. Arthroscopy 2002;18:1013–9. [95] DeRouin A, Pacella N, Zhao C, An K, Ong KG. A wireless sensor for real-time monitoring of tensile force on sutured wound sites. IEEE Trans. Biomed. Eng. 2016;63: 1665–71. [96] Raviglione A, Reif R, Macagno M, Vigano D, Schram J, Armstrong D. Real-time smart textile-based system to monitor pressure offloading of diabetic foot ulcers. J. Diabetes Sci. Technol. 2017;doi: 10.1177/1932296817695339. [97] Lavery LA, Peters EJ, Armstrong DG, Wendel CS, Murdoch DP, Lipsky BA. Risk factors for developing osteomyelitis in patients with diabetic foot wounds. Diabetes Res. Clin. Pract. 2009;83:347–52. [98] Gentzkow GD, Iwasaki SD, Hershon KS, Mengel M, Prendergast JJ, Ricotta JJ, et al. Use of dermagraft, a cultured human dermis, to treat diabetic foot ulcers. Diabetes Care 1996;19:350–4. [99] M. Esposito, M. Macagno, D.G. Vigano. U.S. Patent No. 8925392 B2 (2015).

Additive manufacturing for biofabricated medical device applications

341

[100] Gerhardt L, Lottenbach R, Rossi R, Derler S. Tribological investigation of a functional medical textile with lubricating drug-delivery finishing. Colloids Surf. B: Biointerfaces 2013;108:103–9. [101] Gupta B, Agarwal R, Alam M. Textile-based smart wound dressings. 2010. [102] Massaroni C, Saccomandi P, Schena E. Medical smart textiles based on fiber optic technology: an overview. J. Funct. Biomater. 2015;6:204–21. [103] Jeanne M, Narbonneau F, Witt J, Paquet B, Kinet D, Kreber K . OFSETH: a breathing motions monitoring system for patients under MRI. Conf. Proc. IEEE Eng. Med. Biol. Soc. 2010;2010:1016–9. [104] Ghosh S, Amidei C, Furrow K. Development of a sensor-embedded flexible textile structure for apparel or large area applications. 2005. [105] Quandt BM, Scherer LJ, Boesel LF, Wolf M, Bona G, Rossi RM. Body-monitoring and health supervision by means of optical fiber-based sensing systems in medical textiles. Adv. Healthcare Mater. 2015;4:330–55. [106] Augousti AT, Raza A, Graves M. Design and characterization of a fiber optic respiratory plethysmograph (FORP) 1996;250–7. [107] Deglurkar M, Davy DT, Stewart M, Goldberg VM, Welter JF. Evaluation of machining methods for trabecular metal implants in a rabbit intramedullary osseointegration model. J. Biomed. Mater. Res. B: Appl. Biomater. 2007;80:528–40. [108] Hacking S, Bobyn J, Toh K, Tanzer M, Krygier J. Fibrous tissue ingrowth and attachment to porous tantalum. J. Biomed. Mater. Res. 2000;52:631–8. [109] Kaplan RB. Open cell tantalum structures for cancellous bone implants and cell and tissue receptors 1994. [110] Rodriguez RU, Kemper N, Breathwaite E, Dutta SM, Hsu EL, Hsu WK, et al. Demineralized bone matrix fibers formable as general and custom 3D printed mold-based implants for promoting bone regeneration. Biofabrication 2016;8:035007. [111] Stoeckel D, Bonsignore C, Duda S. A survey of stent designs. Minim. Invasive Ther. Allied Technol. 2002;11:137–47. [112] M.S. Mirizzi. U.S. Patent No. 6,574,851. 10 June 2003. [113] Cabrera MS, Sanders B, Goor OJ, Driessen-Mol A, Oomens CW, Baaijens FP. Computationally designed 3D printed self-expandable polymer stents with biodegradation capacity for minimally invasive heart valve implantation: a proof-of-concept study. 3D Print. Addit. Manuf. 2017;4:19–29. [114] Armstrong EJ, Saeed H, Alvandi B, Singh S, Singh GD, Keong Yeo K, et al. Nitinol self-expanding stents vs. balloon angioplasty for very long femoropopliteal lesions. J. Endovasc. Ther. 2014;21:34–43. [115] Lammer J, Bosiers M, Zeller T, Schillinger M, Boone E, Zaugg MJ, et al. First clinical trial of nitinol self-expanding everolimus-eluting stent implantation for peripheral arterial occlusive disease. J. Vasc. Surg. 2011;54:394–401. [116] Simsekyilmaz S, Liehn EA, Weinandy S, Schreiber F, Megens RT, Theelen W, et al. Targeting in-stent-stenosis with RGD-and CXCL1-coated mini-stents in mice. PLOS ONE 2016;11:e0155829. [117] Warkiani ME, Tay AK, Guan G, Han J. Membrane-less microfiltration using inertial microfluidics. Sci. Rep. 2015;5:11018. [118] Lee WC, Bhagat AAS, Huang S, Van Vliet KJ, Han J, Lim CT. High-throughput cell cycle synchronization using inertial forces in spiral microchannels. Lab Chip 2011;11: 1359–67. [119] Gossett DR, Weaver WM, Mach AJ, Hur SC, Tse HTK, Lee W, et al. Label-free cell separation and sorting in microfluidic systems. Anal. Bioanal. Chem. 2010;397:3249–67.

342

Additive Manufacturing

[120] Pethig R. Review article—dielectrophoresis: status of the theory, technology, and applications. Biomicrofluidics 2010;4:022811. [121] Sharei A, Cho N, Mao S, Jackson E, Poceviciute R, Adamo A, et al. Cell squeezing as a robust, microfluidic intracellular delivery platform. JoVE (J. Vis. Exp.) 2013; e50980–e150980. [122] Golden JP, Justin GA, Nasir M, Ligler FS. Hydrodynamic focusing—a versatile tool. Anal. Bioanal. Chem. 2012;402:325–35. [123] Caves JM, Kumar VA, Wen J, Cui W, Martinez A, Apkarian R, et al. Fibrillogenesis in continuously spun synthetic collagen fiber. J. Biomed. Mater. Res. B: Appl. Biomater. 2010;93:24–38. [124] Haynl C, Hofmann E, Pawar K, Förster S, Scheibel T. Microfluidics-produced collagen fibers show extraordinary mechanical properties. Nano Lett. 2016;16:5917–22. [125] Ferris CJ, Gilmore KG, Wallace GG, Biofabrication: an overview of the approaches used for printing of living cells. Appl. Microbiol. Biotechnol. 2013;97:4243–58. [126] Simon MG, Lee AP. Microfluidic droplet manipulations and their applications. Microdroplet Technology. Springer; 2012. p. 23-50. [127] Toh AG, Wang Z, Yang C, Nguyen N. Engineering microfluidic concentration gradient generators for biological applications. Microfluid. Nanofluid. 2014;16:1–18. [128] Velve-Casquillas G, Le Berre M, Piel M, Tran PT. Microfluidic tools for cell biological research. Nano Today 2010;5:28–47. [129] Judy JW. Microelectromechanical systems (MEMS): fabrication, design and applications. Smart Mater. Struct. 2001;10:1115. [130] A Lifton V, Lifton G, Simon S. Options for additive rapid prototyping methods (3D printing) in MEMS technology. Rapid Prototyp. J. 2014;20:403–12. [131] Whitney J, Sreetharan P, Ma K, Wood R. Pop-up book MEMS. J. Micromech. Microeng. 2011;21:115021. [132] Sreetharan PS, Whitney JP, Strauss MD, Wood RJ. Monolithic fabrication of millimeterscale machines. J. Micromech. Microeng. 2012;22:055027. [133] Gafford JB, Kesner SB, Wood RJ, Walsh CJ. Force-sensing surgical grasper enabled by pop-up book. MEMS 2013;2552–8. [134] Gafford JB, Kesner SB, Wood RJ, Walsh CJ. Microsurgical devices by pop-up book mems. 2013:V06AT07A011-V06AT07A011. [135] Rengier F, Mehndiratta A, von Tengg-Kobligk H, Zechmann CM, Unterhinninghofen R, Kauczor H, et al. 3D printing based on imaging data: review of medical applications. Int. J. Comput. Assist. Radiol. Surg. 2010;5:335–41. [136] Lam CXF, Mo X, Teoh S, Hutmacher D. Scaffold development using 3D printing with a starch-based polymer. Mater. Sci. Eng. C 2002;20:49–56. [137] Majidi C. Soft robotics: a perspective—current trends and prospects for the future. Soft Robot. 2014;1:5–11. [138] Rus D, Tolley MT. Design, fabrication and control of soft robots. Nature 2015;521: 467–75. [139] Bartlett NW, Tolley MT, Overvelde JT, Weaver JC, Mosadegh B, Bertoldi K, et al. SOFT ROBOTICS. A 3D-printed, functionally graded soft robot powered by combustion. Science 2015;349:161–5. [140] Umedachi T, Vikas V, Trimmer BA. Highly deformable 3-d printed soft robot generating inching and crawling locomotions with variable friction legs 2013;4590–5. [141] Mangano F, Chambrone L, van Noort R, Miller C, Hatton P, Mangano C. Direct metal laser sintering titanium dental implants: a review of the current literature. Int. J. Biomater. 2014;2014:461–534.

Additive manufacturing for biofabricated medical device applications

343

[142] De Beer N, Odendaal A. The geometrical accuracy of a custom artificial intervertebral disc implant manufactured using computed tomography and direct metal laser sintering. J. New Gener. Sci. 2012;10:1–13. [143] Jardini A, Larosa M, Bernardes L, Zavaglia C, MacielFilho R. Application of direct metal laser sintering in titanium alloy for cranioplasty 2011;. [144] Salmi M, Tuomi J, Paloheimo K, Björkstrand R, Paloheimo M, Salo J, et al. Patientspecific reconstruction with 3D modeling and DMLS additive manufacturing. Rapid Prototyp. J. 2012;18:209–14. [145] Huang NM. Analyzing the Surface Finish of Knee Implants to Determine Criteria for Applications in Direct Metal Laser Sintering. 2012 NCUR; 2012. [146] Ryan G, Pandit A, Apatsidis DP. Fabrication methods of porous metals for use in orthopaedic applications. Biomaterials 2006;27:2651–70. [147] Matassi F, Botti A, Sirleo L, Carulli C, Innocenti M. Porous metal for orthopedics implants. Clin. Cases Miner. Bone Metab. 2013;10:111–5. [148] Neches RY, Flynn KJ, Zaman L, Tung E, Pudlo N. On the intrinsic sterility of 3D printing. Peer J. 2016;4:e2661. [149] Kolesky DB, Truby RL, Gladman A, Busbee TA, Homan KA, Lewis JA. 3D bioprinting of vascularized, heterogeneous cell-laden tissue constructs. Adv. Mater. 2014;26: 3124–30. [150] Mironov V, Reis N, Derby B. Review: bioprinting: a beginning. Tissue Eng. 2006;12: 631–4. [151] Derby B, Bioprinting: inkjet printing proteins and hybrid cell-containing materials and structures. J. Mater. Chem. 2008;18:5717–21. [152] Guillotin B, Souquet A, Catros S, Duocastella M, Pippenger B, Bellance S, et al. Laser assisted bioprinting of engineered tissue with high cell density and microscale organization. Biomaterials 2010;31:7250–6. [153] Ozbolat IT, Yu Y. Bioprinting toward organ fabrication: challenges and future trends. IEEE Trans. Biomed. Eng. 2013;60:691–9. [154] Submission and Review of Sterility Information in Premarket Notification (510(k)) Submissions for Devices Labeled as Sterile. https://www.fda.gov/downloads/MedicalDevices/../ucm109897.pdf. [155] Technical Considerations for Additive Manufactured Devices Draft Guidance for Industry and Food and Drug Administration Staff. https://www.fda.gov/downloads/MedicalDevices/DeviceRegulationandGuidance/GuidanceDocuments/UCM499809.pdf. [156] Consider the Basics when Developing Packaging. http://www.mddionline.com/article/ consider-basics-when-developing-packaging. [157] Guideline for Disinfection and Sterilization in Healthcare Facilities https://www.cdc. gov/infectioncontrol/pdf/guidelines/disinfection-guidelines.pdf (2008). [158] Dai Z, Ronholm J, Tian Y, Sethi B, Cao X. Sterilization techniques for biodegradable scaffolds in tissue engineering applications. J. Tissue Eng. 2016;7. 2041731416648810. [159] Augustine R, Saha A, Jayachandran V, Thomas S, Kalarikkal N. Dose-dependent effects of gamma irradiation on the materials properties and cell proliferation of electrospun polycaprolactone tissue engineering scaffolds. Int. J. Polym. Mater. Polym. Biomater. 2015;64:526–33. [160] Lee JB, Ko Y, Cho D, Park WH, Kim BN, Lee BC, et al. Modification of PLGA nanofibrous mats by electron beam irradiation for soft tissue regeneration. J. Nanomater. 2015;16:136. [161] Hofmann S, Stok KS, Kohler T, Meinel AJ, Müller R. Effect of sterilization on structural and material properties of 3-D silk fibroin scaffolds. Acta Biomater. 2014;10:308–17.

344

Additive Manufacturing

[162] Stephens B, Azimi P, El Orch Z, Ramos T. Ultrafine particle emissions from desktop 3D printers. Atmos. Environ. 2013;79:334–9. [163] Nanofibers used in first-ever synthetic tracheal implant. http://medicaldesign.com/materials/nanofibers-used-first-ever-synthetic-tracheal-implant. [164] Sachs E, Cima M, Cornie J, Brancazio D, Bredt J, Curodeau A, et al. Three-dimensional printing: the physics and implications of additive manufacturing. CIRP Ann. – Manuf. Technol. 1993;42:257–60. [165] Sachs EM, Haggerty JS, Cima MJ, Williams PA. Three-dimensional printing techniques 1993. [166] Deckard CR. Method and apparatus for producing parts by selective sintering 1989. [167] Ackland DC, Robinson D, Redhead M, Lee PVS, Moskaljuk A, Dimitroulis G. A personalized 3D-printed prosthetic joint replacement for the human temporomandibular joint: from implant design to implantation. J. Mech. Behav. Biomed. Mater. 2017;69:404–11. [168] Chandak P, Byrne N, Newton V, Coleman A, Stojanovic J, Marks SD, et al. Classification of abdominal vascular anomalies and use of 3D printing to support complex renal transplantation in children. Lancet 2017;S32. [169] Crump SS. Apparatus and method for creating three-dimensional objects 1992.