To infinity and beyond: the promise of data-driven 3D printing of hernia mesh – a primer for surgeons

3D printing is a form of additive manufacturing that has been in existence for some time. Raw materials suitable for 3D printing include metals, ceramics, paper and polymers. Using computer-aided design (CAD) software, a desirable object is digitally created and saved as a CAD file. CAD files may also be created by scanning objects using specialised laser equipment, or by reconstructing objects from radiological imaging, such as Digital Imaging and Communications in Medicine (DICOM) files. Using CAD files, segmentation software then digitally slices the object into thin layers and creates a printable instruction file known as G-code. G-code is then transmitted to 3D printers, and the object is created [37].

Many types of 3D printing are available on the market, with the difference primarily being how the raw material is prepared and how it is bound together to form the object. Some 3D printers, such as electron beam melting (EBM) or direct energy deposition (DED) are only suitable for the creation of metallic or ceramic objects, typically used in automobile industry or aerospace engineering. The following 3D printers are some examples with relevance to producing custom hernia mesh.

Fused filament fabrication (FFF), also known as Fused Deposition Modelling (FDM), is one of the initial 3D printers developed and was first patented in 1989 [38]. FDM has experienced rapid growth in the technology in recent decades [39], primarily due to use in rapid prototyping [40]. Rapid prototyping is the designing and printing models of an object or feature within a short time frame, often as a sample (prototype) to a bigger project. In FDM, once G-code has been transmitted to the 3D printer, polymer filaments are heated to a liquid state, extruded through a nozzle and deposited layer by layer. Specific heating and nozzle settings for a given polymer material are supplied by filament manufacturers [41]. FDM is considered user-friendly and has a low setup cost. It can use a range of common thermoplastics at relatively low temperatures, such as polylactic acid (PLA), acrylonitrile butadiene styrene (ABS), polyethylene terephthalate glycol (PETG) or thermoplastic polyurethane (TPU) [42]. High-end FDM printers can achieve greater temperatures, allowing for the use of high-performance polymers such as polyether ether ketone (PEEK), polyetherimide (PEI) or carbon fibre composites [43]. FDM is versatile, and a good introduction to 3D printing for novices. Drawbacks of FDM include loss of dimensional accuracy during printing, printing may be time consuming for complex geometry, and post-processing (e.g. sandpapering or milling) may be required to achieve the desired surface finish.

Stereolithography (SLA) uses an ultraviolet (UV) laser to cure liquid resin layer by layer, allowing finer detail and smoother prints to be achieved [44]. Objects generally require post-processing, including curing in a UV chamber to further harden the object [45], and application of isopropyl alcohol (IPA) to remove surface tackiness [46]. 3D-printed resin models are commonly used for teaching purposes, to demonstrate patient-specific anatomy or for pre-operative planning. Most resins are epoxy-based with carcinogenic properties, and thus are considered too toxic for implant applications [47, 48].

Selective laser sintering (SLS) uses lasers to fuse powder particles layer by layer. Powder can be metallic, nylon or TPU. Without the need for support structures, complex geometries and intricate internal structures may be created [49, 50]. SLS is well suited to produce hernia mesh and surgical-grade implants. SLS is currently only available in industrial-size platforms, and initial set-up costs may be considered prohibitive [51].

Bioprinting is a fusion of 3D printing technology and tissue engineering [52]. Instead of using inorganic materials, bioprinting uses bioink. Bioink is a water-based hydrogel that mimics the extracellular matrix, and typically incorporates cells and biochemicals, such as growth factors and cytokines, to support vascularisation and tissue growth [53‐56]. Bioprinting have been used experimentally to produce custom organs, and has immense potential in regenerative and transplant medicine [57]. Bioprinting could theoretically be used to create custom biological hernia meshes with bioactive properties that elicits desirable immunological responses to initiate healing, with minimal foreign body response.

Commonly used medical-grade polymers include polycaprolactone (PLC) [58‐66], polylactic acid (PLA) [67‐69], polyvinyl alcohol (PVA) [70, 71], polypropylene (PP) [70, 72], polyurethane (PU) [2, 73] and gelatine carbon nanotubes [74]. These polymers have minimal toxicity, are well established in the medical industry, and can be easily converted into printable 3D filament with a filament extruder.

3D printing has created an alternate pathway towards creation of affordable custom biomedical devices, a feat previously not possible with traditional manufacturing techniques. The high degree of customisation offered by CAD allows designing and manufacturing processes to be driven predominantly by clinician expertise and patient circumstances, creating a collaborative environment to produce patient-specific implants and tailored healthcare. Patient-specific implants for orthopaedic, dental and facial reconstruction have been shown to achieve faster functional return and greater patient satisfaction [75‐78]. Detailed anatomical models have been invaluable to surgical education and simulations [79‐81].

In the case of abdominal wall hernia management, a range of customisable features may be possible [59‐74, 82, 83]. The size, contour and shape of the mesh could be predetermined based on preoperative imaging, using either computed tomography (CT) or magnetic resonance imaging (MRI) [84]. This could remove the need to trim meshes intraoperatively, decrease handling time and reduce risk of bacterial contamination. Optimal filament size and mesh thickness may be predetermined based on expected wall stresses and desired effective porosity. The thickness of the mesh may be varied such that it has adequate strength over maximal stress regions, as dictated by fracture mechanics, while minimising mesh burden [18]. Specific mesh material may be chosen based on expected activity levels [84]. For example a flexible bioabsorbable polymer may be better suited for a hernia repair in a young patient who is expected to lead a more active lifestyle.

Mesh surface topography can be laser etched to dissuade bacterial colony attachment and subsequent biofilm formation [85]. Use of plasma processing could increase surface hydrophilic properties to encourage cell attachment [86, 87]. Radiopaque contrast may be incorporated into mesh fibres to allow better visualisation on subsequent medical imaging for follow-up or diagnostic purposes [59]. Likewise, drug release microcapsules, bacteriophages, zinc, copper or silver ions could be incorporated into the mesh structure to provide antibacterial properties, and potentially reduce post-operative infections [62, 64, 67, 70, 73, 74, 88, 89]. A new generation of ‘intelligent’ meshes may be possible by imbedding special sensors in the mesh structure to detect infection or monitor healing [90].

The potential of 3D-printed meshes may also be extended to a broader range of indications where meshes have been conventionally used. Pelvic floor reconstructions are notoriously difficult to perform due to geometric shape, dynamic structures, and prominent neurovascular bundles [91]. Patient-specific 3D-printed implants could reduce the difficult of such operations, and facilitate local tissue integration with adjuncts built into the mesh structure. Similarly, congenital hernias in paediatric patients require accommodation of body growth when planning surgical repairs, and 3D-printed implants could offer improved outcomes [92].

The utility of 3D-printed patient-specific implants is likely applicable to all surgical disciplines that require individualised anatomically responsive solutions. With sufficient trained personnel and equipment, patient-specific implants could be produced onsite within a healthcare facility [19, 93]. 3D printing is an attractive consideration by centres in developing regions where normal procurement may be difficult due to logistic supply chain issues [94].

Clinical use of 3D-printed hernia mesh is currently limited to case-by-case provision [84]. The vast majority of published papers are preclinical and focus on describing mesh production techniques, in vitro toxicity, performance of adjuncts (e.g. antimicrobial properties) and short-term tolerance by small mammals [60, 62‐64, 68‐71, 82]. Only one study used a large animal model, namely sheep [65]. While there is nothing inherently wrong with using small animal models, and perhaps encouraged from an animal welfare point of view when investigating new substances and products [95], in vivo outcomes observed in small animals typically do not translate well to performance in large animals or humans [96]. Large animal studies over medium- to long-term are necessary to characterise in vivo behaviour of 3D-printed mesh products, as mandated by the International Organization for Standardization 10993-6 [97].

The meshes printed so far are relatively small. Of the studies that reported mesh size, only two studies implanted meshes larger than 5 × 5 cm [70, 82]. Uniaxial tensile testing of printed mesh was reported in 15 out of 19 experimental studies. Of which, only 3 studies compared their tensile strength to the estimated longitudinal stress of 16 N/cm and hoop stress of 32 N/cm in the abdominal wall, as derived by Klinge et al. [98] Works by Kallinowski’s group indicate intra-abdominal pressures during intense post-operative vomiting may be as high as 225 mmHg [99]. This rapid change in pressure could theoretically generate a hoop stress of 48 N/cm² along the abdominal wall, or about 48 N/cm for every 1 cm length of tissue [100]. The 3D-printed meshes of some studies are likely to fracture before they can obtain the necessary GRIP values to overcome the CRIP values of defects [24]. Coined by Kallinowski, GRIP is a numerical representation of the stability of mesh fixation and surface friction it has to resist against cyclical impacts, while CRIP describes the minimum level of GRIP required to resist against a standardised quantity of cyclical impacts that simulates intra-abdominal pressures generated by post-operative coughing. On the other hand, some meshes are likely overengineered, such as Erwin et al.’s 3D-printed polypropylene mesh that has a reported tensile strength of 321.67 kgf/mm² (~ 315,000 N/cm²) which is on par with Kevlar fibres (362,000 N/cm²) used in ballistic body armour [101].

Elasticity is another consideration and appears to be poorly documented. Only two studies clearly reported elongation of 3D-printed mesh under strain [68, 73]. The abdominal wall has an inherent degree of elasticity (about 32% at 16 N/cm) which the mesh should ideally match [102]. This is influenced by the construct of the mesh, with lightweight mesh (35–70 g/m²) stretching more than heavyweight mesh (≥ 140 g/m²) [103, 104]. Meshes with elongation rates greater than the native abdominal wall, i.e. >30%, may not maintain functional repair and could be a cause for repair failure in the long-term [100]. Excess elasticity can also alter effective porosity [105]. While several studies reported textile porosity over 1 mm² [60, 62, 68, 70, 73, 74, 106], which is the minimum pore size needed to prevent the bridging effect from fibrotic tissue [107], no study specified whether these values were under static or dynamic conditions. Loss of effective porosity prevents ingrowth of tissue and leads to excessive scar plate formation, which may precipitate development of seroma, chronic pain and repair failure [11, 13].

Another problem is long-term stability of mesh. The most commonly reported materials used to fabricate 3D-printed meshes were PCL and PLA (Table 4). Both of these substances have well-established biological safety and immune profiles [33]. PCL is a copolymer of the monofilament suture Monocryl^® (polycaprolactone/polyglycolide) (Ethicon Inc, New Jersey, United States), while PLA is the primary component of the ‘Velcro-like’ grips in ProGrip^® mesh (Medtronic Australia Pty Ltd, New South Wales, Australia). Both PCL and PLA are classified as biodegradable polymers. These substances are currently only used as adjuncts or as a copolymer in regulatory-approved hernia mesh, and not in their pure forms. Although PCL is reported to have a ‘slow degradation’ over 2–4 years [108, 109], concerns have been raised that electrospun fibres have accelerated degradation when exposed to hydrolytic enzymatic action, and tensile strength loss may be encountered as early as 90 days [110]. The manufacturer data on Monocryl^® indicates an expected loss of strength over 14 days, and complete resorption by 90–120 days. Investigations into using PLA as a suture material found 12% reduction in knot strength after submersion in normal saline at room temperature for 28 days [111]. A polymer mixture or composite structure may be required to meet the biomechanical demands of hernia repairs, while still having low immunogenicity and biodegradative properties.

Since hernia meshes are long-term implantable devices, sterility of the mesh must be guaranteed. It is suspected that bacterial contamination, such as by Staphylococcus aureus [112], and subsequent biofilm formation from sub-acute infection is responsible for a portion of repair failures years after surgery [113]. In most of the experimental papers, if described, have only used submersion in concentrated ethanol and UV light exposure as a sterilisation method (Tables 4 and 5). These methods should be more appropriately termed disinfection and are inadequate for sterilising critical devices, such as surgical implants, that enter sterile regions of the body, due to their inability to kill and remove bacterial spores [114, 115]. A sterilisation study of 3D-printed objects noted that hollowed objects could not be fully sterilised with either steam autoclave or ethylene oxide (EO), with non-coagulase Staphylococcus species continued to be isolated post-sterilisation [116]. Likewise, Garnica-Bohorquez et al. did not find any success in a two-step formaldehyde steam autoclave and noted significant loss in tensile strength of 3D-printed PLA meshes. Shea et al. tested low-temperature vaporised hydrogen peroxide (VHP) gas plasma sterilisation for a variety of 3D-printed objects for clinical use, and recorded a surgical site infection rate of 7.0% (8 of 114 patients), with 5 patients (62.5%) requiring surgical debridement or implant removal/revision [117]. Although the temperature of printing processes typically exceeds 200° Celsius, a substantial amount of post-printing processing is expected and opens up multiple routes and opportunities for contamination [117]. The surface topography of 3D-printed objects is irregular, and like textured breast implants or healthcare surfaces, provides the ideal environment for bacterial attachment and biofilm formation [118, 119].

EO and VHP are standard methods of sterilisation in the central sterilisation services department (CSSD) in hospitals, with typical operating temperatures of 50–60° Celsius. This is close or substantial above the glass-transition temperatures of PLA or PCL, which are 60–65° Celsius and minus 60° Celsius, respectively [120]. Sterilising with EO or VHP using the current protocols will likely lead to deformation and surface damage of 3D-printed objects [120]. Currently, it appears STERIS Healthcare (Dublin, Ireland) is the only company providing a dedicated sterilisation cabinet for 3D-printed objects using VHP (V-Pro Max 2 system), in conjunction with proprietary resin materials. It is unclear whether such resin materials are suitable for hernia mesh production. Choice of mesh printing material needs to take into consideration of method of sterilisation, and the potential physiochemical reactions that may occur.

Without large animal model and clinical studies, it is difficult to state whether 3D-printed mesh will be efficacious to the management of abdominal wall hernias. The experiments identified so far indicates that there is no immediate toxicity or handling concerns. Histological analysis consistently shows expected acute phase inflammatory reaction [60, 62‐65, 68, 70, 71, 82, 106]. No further meaningful conclusions could be drawn due to high risk of bias from unclear randomisation process of animals involved (Fig. 2). To maximise data reliability, the PREPARE and ARRIVE guidelines should be followed when designing and reporting animal studies, and the equivalent guidelines for in vitro studies [32, 121]. Pre-registration of animal protocols is strongly recommended to ensure study integrity and minimisation of biases.

A critical element in evaluating whether a hernia mesh function as intended, is to assess the interface between the mesh and tissue, and to examine how well tissue integrates or infiltrates over time. Good tissue integration is a reflection of beneficial tissue growth and minimal foreign body response, and corresponds with strong biomechanical stability that is necessary at preventing long-term hernia recurrence. To assess mesh tissue integration in an objective fashion, a standardised Mesh Integration Index was previously proposed and validated in a large animal model [manuscript under consideration] [122]. The Index allows preclinical standardised assessment of mesh performance, providing the necessary information to inform researchers whether a mesh function as intended. Problematic meshes could be identified and recalled prior to marketing, and thereby minimising harm to patients. The Index is applicable to all mesh products in the abdominal wall, including 3D-printed meshes.

Briefly, the Index uses a series of standardised assessments to grade integration, fibrosis, degradation and adhesion on a 0 to 5 ratio scale. The assessments incorporate visual, histological, biomechanical and molecular tools. These tools were selected due to their widely available in standard biomedical research institutes or laboratories, which allows rapid assessment of in vivo results using standard animal models. The in-vivo behaviour of 3D-printed meshes is easily quantified by the Index, streamlining comparison of different 3D-printer compounds and discovery of new mixtures or methods to produce clinically impactful hernia meshes. The secondary purpose of the Index is to provide a uniform language between biomaterial scientists, who are likely involved in the designing of the product, and clinicians, who are ultimately the end users of mesh products.

The ideal mesh is one that achieves high levels of tissue integration in the shortest amount of time after implantation and has the least amount of foreign body fibrotic reaction. The mesh should conform to the structure of the local anatomical structures, have sufficient elasticity, maximal effective porosity, antibacterial properties and sufficient strength to accommodate the dynamic physiological stresses in the abdominal wall. Meshes that intend to be intraperitoneal should have anti-adhesion properties that do not interfere with tissue integration. Meshes should ideally have minimal degradation with no loss of tensile strength for at least 5 years from time of implantation. Meshes should preferably be relatively easy to handle, can tolerate intraoperative manipulations, conform to local anatomy and not require trimming.

Like any other medical device, 3D-printed mesh, if deployed to clinical use, should be tracked and monitored with a clinical quality registry that incorporates a device registry. Outcomes of lesser-known products should be monitored, such that adverse events can be detected at the first instance, to minimise harm to patients. Registries and big data platforms are vital to tracking real‑world effectiveness and complications (e.g. China’s national hernia registry with ~ 100,000 cases) [123]. Ultimately, device regulations should occur at a national level, through regulatory authorities, such as the United States Food and Drug Administration (FDA), the European Union (EU), or the Australian Therapeutic Goods and Administration (TGA).

Medical device regulation has traditionally been manufacturer heavy, with the onus on only approving devices that meet the necessary standards for safety and efficacy as set by the local regulatory authority. Devices undergo a series of standardised testing to ensure that that it is not toxic to humans, it achieves its intended purpose, and it does not have long-term side effects or potential problems that may arise from device malfunction. Devices need to be manufactured to a minimum standard that meets national and international regulations, ensuring integrity and consistency of the device, be adequately sterile for its purposes, and have quality assurance processes in place to detect problems. 3D-printed hernia meshes currently falls into a grey area within legislation, as it can be both mass-produced in terms of numbers and has a degree of customisation to fit the patient for usage [124, 125].

In the US, oversight relies on the FDA’s Custom Device Exemption (CDE), such as Investigational Device Exemption (IDE) via Centre for Devices and Radiological Health (CDRH) or the 510(k) Premarket Approval (PMA) pathway. The latter is optimised for conventional ‘substantially equivalent’ devices. Under the CDE pathway, a 3D-printed device may qualify as exempt from standard premarket approval if it meets specific criteria: it must be custom-made in direct response to a written request from an authorised health professional (e.g. a surgeon); the design must originate from the clinician and not be substantially manufactured in advance by the device producer; often, design input may include patient-specific templates, scans, or clinical sketches provided by the health professional [126]. Truly personalised meshes may struggle to fit existing categories but if the mesh is fully unique per patient (e.g. from a CT scan-derived 3D model) and fabricated as a once off, CDE may apply but this cannot be used for scaling up or commercialisation under the 21 Code of Federal Regulation (CFR) 812.3(b). If a 3D-printed mesh has no predicate but poses low to moderate risk, de novo classification may be achieved instead via US Federal Food Drug and Cosmetics Act (FDCA) Sect. 513(f)(2) [127]. Dykema argues that a new scheme should be created by the FDA under the existing Class III medical devices regulatory classification that is specific to custom 3D-printed devices and contains elements designed specifically to address the unique nature of additive manufacturing at point of care [124].

In the EU, 3D‑printed devices fall under the Medical Device Regulation (MDR). The EU MDR (2017/745/European Union) outlines the framework for regulation, mandating a bare minimum of clinical evaluation, risk management, quality management system, post-market surveillance, technical documentation, and liability for defective devices [128]. A submission for use to MDR requires detailed technical documentation and clinical justification equivalent to major implants. This places significant demands on healthcare facilities to coordinate multidisciplinary in‑house teams, certified under International Organization of Standardization (ISO) 13,485, if 3D printing was to be performed at point of care [129]. Pettersson critiques that the MDR does not truly address hospitals as manufacturers, and there continue to lie problems of ownership and usage of intellectual properties of devices and software, not withstanding liability issues [129].

In Australia, surgical meshes were reclassified by the TGA from Class IIb (medium risk) to Class III (high risk) devices in 2018, and fully implemented by 2021 [130]. These changes mandated a full TGA conformity assessment and inclusion within the Australian Register of Therapeutic Goods (ARTG), removing the previous lenient pathway for implant approval [131, 132]. The TGA currently does not have a dedicated pathway, and references the technical considerations set up by the US FDA and the local Therapeutic Goods (Medical Devices) Regulations 2002 [133].

In addition to the regulatory difficulties, use of 3D-printed mesh has ethical ramifications for service provision and duty of care. To produce 3D-printed mesh, a template is developed based on patient medical imaging data, and the product is created inside a designated 3D printer, using a pre-specified compound [128, 134]. Each of these step may require a proprietary substance, product or service, and involvement of multidisciplinary personnels. Sharing of patient medical data may be required, and the potential for data breach may have serious consequences. These aspects may not necessarily be fully apparent to clinicians, yet they are expected to provide full disclosure to patients and be ethical gatekeepers. Recent medical device litigation, such as the US$1 billion mesh settlement by Becton Dickinson & Co in 2024 [135], have illustrated the need for transparency in medical products and clinical services to which biomedical companies have a vested financial interest. Although clinicians may have the best intentions for patients, use of incomplete or incorrect information may inadvertently lead to harm. Likewise, clinicians should also not be the sole provision of 3D-printing services, as the potential for misconduct and harm to patients to occur in a clinical setting without oversight is high, as illustrated by Foster’s case studies [125].

The safest option, is perhaps a middle ground, where 3D-printing services are provided onsite at the healthcare facility via service agreement with a certified manufacturer. Service agreements could be established to provide a comprehensive 3D printing service package, that incorporates provision of printers, materials and trained technicians. This could be an option in minimising the legal issues of ownership of proprietary knowledge and liability of healthcare facilities during manufacturing, quality control and sterilisation of 3D-printed meshes. Responsibility is distributed to key individuals, namely physicians for selection of products; radiologists and technicians for medical imaging scanning; technicians for segmentation of imaging for 3D printing; engineers for product designing and 3D printing; and physicians and sterilisation units for the final preparation, sterilisation and usage of product [128, 129]. Such business models are not new to the healthcare system, and much could be learned by examining the precedent set by the local blood transfusion services, such as the Northern Ireland Blood Transfusion Service and the Australian Red Cross Life Blood^®. Both services manufacture blood products from patient blood donations under highly regulated yet transparent systems, and ensure all products are accountable, traceable and reportable in the event of adverse reactions.

Patient safety should be foremost when dealing with emerging technology and implantable devices. All 3D-printed hernia mesh should be produced to a high quality that meets national and international standards. An oversight committee, run independently by the national regulatory authority, could provide service accreditation and maintain a robust pre- and post-market surveillance of safety and efficacy. Sharing of patient information, such as DICOM files, should adhere to local privacy policies and data protection requirements, such as the General Data Protection Regulation (GDPR) for nations in the European Union. An encrypted collaborative platform on a local area network with time-limited and personnel-limited access may be required to ensure no unnecessary retention of patient information during the manufacturing process. A designated data compliance team could regulate access and monitor for potential data breach. 3D printing should also take into consideration of waste production, toxicity of waste and environmental impact of polymers and microplastics. Waste should be disposed of in a safe and regulated manner, as dictated by local legislations and international laws.

The true cost of 3D-printed hernia mesh is unknown at this stage, and much of it is considered proprietary knowledge by companies who are attempting to establish themselves within the 3D-printed patient-specific implant market. The first FDA 510(k) clearance for a 3D-printed surgical mesh was only granted in May 2024, awarded to the US-based PrintBio, Inc.’s 3DMatrix™ product (K232602) [136]. Made from polydioxanone monofilament, it has a macroporous non-woven fully-absorbable architecture that has been theorised to reduce complications and healthcare costs in the long term [73]. More information is expected to come to light as products and services of respective companies are approved by the FDA, or equivalent authority.

While functionally not entirely the same, some insights can be gained by examining 3D-printed orthopaedic implants, which have been in use for some years now. The global market for 3D-printed orthopaedic implants was valued at US$1.7 billion in 2024, and is projected to increase to $6.6 billion by 2033 [137]. In Belgium, 3D-printed hip implants have been reported to cost around €8419, comparable to traditional custom-made implants cost of €6002 [138]. Use of 3D-printed hip implants in revisions surgery, compared to traditional implants, can lead to an annual saving of €1,265 with a 5% gain in quality-adjusted life years (QALY) [138]. Similarly, 3D-printed polyether ether ketone (PEEK) cranial implants are reported to cost around US$5,600 to US$20,522 [139]. While such PEEK implants are about 20–133% more expensive than traditional titanium or polymethyl methacrylate (PMMA) implants, they can achieve an excellent successful implantation rate of 93.7%, with minimal complications [139].

A hypothetical business model would suggest an initial capital investment for setup, including equipment, materials, engineering staff, and an ongoing cost for maintenance and consumables. Such a model would likely not be significantly different from implementation of robotic surgical platforms [140]. Once in-house systems are operational, repeatable workflows could reduce per-unit cost over time [141]. The raw materials or consumables required for 3D-printed implant product have been reported to be relatively low [142]. If the upfront costs can be offset by the long-term reductions in complications and reoperations, a viable business model may be possible. In the US, hernia recurrence and associated complications currently account for 20% of annual health expenditure spent on managing incisional hernias in both inpatient and outpatient settings [143].

Health economic assessments, such as cost-benefit analysis and cost-effectiveness analysis will need to be performed as more information emerges, particularly product details, clinical performance, and patient outcomes. Incorporating long-term cost-offsets into cost-effectiveness analyses could pave the way for organisational reimbursement, and overcome cost-related adoption barriers faced by healthcare facilities [144].

3D-printed hernia meshes promises to be an exciting development in tailored hernia care. The potential to use custom-made hernia mesh tailored to individual circumstances, both anatomical shape and tissue environment, could provide an alterative to current therapeutic options. Smart meshes with inbuilt antimicrobial resistance could be the next-generation implants that address the various implantation problems encountered by current clinicians and may improve long-term patient outcomes. The exact parameters of 3D-printed meshes will need to be fine-tuned, and a partnership between surgeons and biomedical engineers at institutions with a strong biomedical engineering program is likely a good foundation to begin with. Early involvement of regulatory bodies and biomedical companies with an interest in developing and providing 3D-printing services at point of care could address the many issues identified in this review.

Likewise, safety and general in-vivo behaviour of 3D-printed meshes should be characterised preclinically, using standardised in-vivo animal models that mimic the human body and facilitate objective comparison, such as the Mesh Integration Index. Long-term outcomes of 3D-printed meshes should be monitored via a clinical quality registry (CQR), as the typical time frame for mesh complications and failure is in the order of 5 years at a minimum [9]. A registry randomised control trial (RRCT), using data from a well-maintained CQR, is an emerging and acceptable alternative to traditional randomised control trials (RCT) in assessing real-world implementation of interventions, particularly over a long period of time that is often not feasible for RCTs due to running costs [145].

Hernias, as a disease entity, is complex, and a universal solution for all scenarios is likely not possible [17]. In the current age of evidence-based medicine, data-driven decision-making is becoming increasingly important in delivering individualised care to patients [146]. There is increasing evidence that hernia care is entering a new phase focused on individualised care and clinicians should strive for data-driven bespoke management.