Bioprinted Constructs Incorporating Mesenchymal Stromal Cells for Osteochondral Repair

Vukovic D¹, Molina C², Gallagher C³

Introduction

The osteochondral unit is a graded composite in which hyaline cartilage is anchored, through a calcified interface, to subchondral bone. Damage rarely respects that boundary, and the surgical options that exist repair neither compartment convincingly. The problem was already stated plainly two decades ago: cartilage repair procedures generate tissue whose composition and mechanics differ from the native articular surface, and durable restoration of a load-bearing joint surface has not been demonstrated [1]. Mesenchymal stromal cells (MSCs) offer a single autologous cell source capable of both chondrogenic and osteogenic differentiation, and three-dimensional printing offers a way to place them, with distinct matrices, in the two compartments at once. The combination has been pursued for more than a decade, beginning with printed topologically defined cell-laden hydrogel scaffolds intended explicitly as osteochondral tissue equivalents [2].

The field has developed a coherent vocabulary and a mature methodological base. Bioinks are now formally distinguished from biomaterial inks by whether cells are present at the moment of processing [3]; extrusion has become the dominant deposition route and its process parameters are well characterised [4]; the polymer chemistry available for cell-laden printing has been reviewed and extended repeatedly [5]. Support-bath printing removed the requirement that a bioink support its own weight during deposition [6], tandem thermal and photo-triggered gelation decoupled the rheology needed for extrusion from the crosslink density needed afterwards [7], and cartilage-derived progenitor cells have been shown to outperform expanded chondrocytes in printable hydrogels [8].

What has not developed is mechanical competence. A printed MSC-laden construct leaves the printer at roughly 5 kPa and, after eight weeks of chondrogenic culture, reaches roughly 42 kPa [9]. Native articular cartilage is quoted at 400-800 kPa in bulk compression [10], and human subchondral bone at 13.34 +/- 1.69 GPa by nanoindentation [11]. This review is organised around that gap rather than around the printing modalities, because the gap is the reason no printed MSC construct has entered clinical use, and because most of what the field reports as progress is measured on axes that do not close it.

The Mechanical Gap

The numbers must be handled with care, because the literature routinely places quantities of different physical type on the same page. Three distinct measurements circulate. Bulk unconfined compression gives an apparent modulus of a whole construct, dominated by hydrogel water content and by any fibre network present. Dynamic nanoindentation gives a shear storage modulus at tissue scale. Quasi-static nanoindentation gives an elastic modulus of mineralised matrix at the micrometre scale. These are not interchangeable, and a comparison that pools them overstates or understates the gap depending on which direction the error runs.

Within bulk compression, the record for cell-laden printed constructs is narrow. MSC-laden norbornene-modified hyaluronic acid printed with in-capillary photocrosslinking measures 5.2 +/- 1.5 kPa as printed and 42.0 +/- 13.9 kPa after 56 days of chondrogenic culture [9]. Unreinforced hydrogels of the kind used across the field sit in the same decade: gelatin-methacryloyl at 7.1 +/- 0.5 to 7.5 +/- 1.0 kPa and alginate at 15.8 +/- 2.0 kPa in compression [10]. The only cell-laden printed construct approaching native cartilage in compression is an inkjet-printed PEGDMA scaffold at 395.73 +/- 80.40 kPa as printed [12], a value achieved through crosslink density rather than through deposited matrix, and one that says nothing about what the encapsulated cells will build.

On the bone side the disparity is categorical rather than incremental. In twelve human cadaveric knees aged 31 to 88, quasi-static nanoindentation gave femoral subchondral bone 13.34 +/- 1.69 GPa and tibial subchondral bone 13.46 +/- 1.78 GPa, with trabecular bone at 12.65 +/- 1.79 GPa and 12.10 +/- 2.36 GPa respectively; articular cartilage in the same specimens gave a shear storage modulus of 0.14-1.30 MPa by dynamic nanoindentation [11]. Between a day-zero MSC bioink and femoral subchondral bone lies something over six orders of magnitude, and even the stiffest reinforced construct remains roughly four and a half log units short. The honest framing is that the chondral compartment is within about two log units of its target at implantation and about one after two months in vitro, while the bone compartment is not being addressed by hydrogel

modulus from 12 human cadaveric knees (ages 31-88) and is not directly comparable with them [9,12,10,11].

Printing Modality and the Viability Record

Against that mechanical picture, viability is the metric the field reports most and interprets least carefully. Values are quoted at day 0, day 1, day 7 and day 21, by live/dead imaging and by metabolic assay, and occasionally as a relative change with no absolute denominator. Extrusion of MSCs in agarose, alginate, GelMA and a commercial bioink gave approximately 80% viability post-printing across all four, assessed by live/dead before 28 days of TGF-beta3 chondrogenic culture [13]. Human chondrocytes in nanofibrillated cellulose-alginate gave 73% at day 1 rising to 86% at day 7 [14]. Human bone-marrow MSCs printed by coaxial extrusion in GelMA with chondroitin sulfate-AEMA and methacrylated hyaluronic acid held 85-90% across assessments at 7, 14 and 21 days, at cell densities above 1 x 10^7 cells/mL and a printing resolution near 100 um [15]; that study reports in vitro neocartilage formation and is frequently mis-cited as an osteochondral regeneration study, which it is not. In-capillary photocrosslinking through a 0.8 mm capillary sustained above 85% at day 7 with 750-820 cells/mm^2 through the construct depth and approximately 3% dimensional error [9].

Inkjet deposition performs comparably. Human bone-marrow MSCs in acrylated-peptide/PEG hydrogel with simultaneous photopolymerisation gave 87.9 +/- 5.3%, though the assay timepoint is not stated [16], and hybrid inkjet printing alternated with PCL electrospinning gave above 80% at one week in five-layer constructs 1 mm thick [17]. Photopolymerising PEGDMA during rather than after printing raised chondrocyte viability by 26% [12] — a relative figure that cannot be converted into an absolute one. Tomographic volumetric printing, which forms a centimetre-scale construct in seconds rather than by serial deposition, reports above 85% [18] and above 90% after a 30 s print of hMSCs with HUVECs in 5% w/v GelMA with 0.05% LAP, with osteogenic markers followed to 42 days and a significant endothelial co-culture effect at 21 days [19]. Laser-induced forward transfer of MSCs is reported to cause no detectable cell damage and to permit densities sufficient for chondrogenesis, but no numeric viability or resolution value is available for it [20], leaving a genuine hole in any modality comparison.

Two observations follow. First, the spread across modalities is smaller than the spread within extrusion, where the same technique yields 46.5 +/- 1.3% at a gel-ceramic interface on day 1 and above 95% on day 0 elsewhere [21,22]. Second, viability is a poor discriminator because nearly every reported figure clears the threshold at which the construct's failure mode becomes mechanical rather than biological. The process variable that actually constrains extrusion is shear stress, which trades directly against resolution and stem cell integrity [23].

Fibre Reinforcement and the Osteochondral Interface

Because bioink chemistry alone cannot reach cartilage-level stiffness without crosslink densities that suppress matrix production, reinforcement has become the principal structural strategy. Melt-electrowritten PCL microfibre networks are themselves compliant — 1.1 +/- 0.3 kPa at 98% porosity and 15.2 +/- 2.2 kPa at 93% porosity in compression — yet infiltrating the 93%-porous network with GelMA yields 214 +/- 24 to 405 +/- 68 kPa depending on initiator concentration, a synergistic increase of up to 54-fold over either component [10]. That places a composite within the 400-800 kPa band quoted for native cartilage, and it is the clearest demonstration that the gap is closable by architecture rather than by chemistry. Cartilage-ECM-functionalised bioinks combined with printed fibre networks pursue the same route while retaining a matrix that supports MSC chondrogenesis [24].

showing a synergistic increase of up to 54-fold [10].

Reinforcement does not, however, transfer to the cellular compartment for free. The composite values above are properties of the construct, not of the deposited matrix, and a fibre network that carries load also shields the encapsulated cells from the strains that drive chondrogenic matrix production. No study in this set reports a cell-laden reinforced construct measured in compression at both day 0 and a late timepoint, which is precisely the comparison that would show whether reinforcement and biological maturation are additive or antagonistic.

The interface between compartments is a separate problem. Interlocking melt-electrowritten meshes protruding from a printed alpha-TCP/nano-hydroxyapatite bone compartment raised hydrogel-to-ceramic adhesion strength more than 6.5-fold relative to non-interlocking architectures and gave the chondral compartment a 20-fold reinforcement over the pristine hydrogel, with newly synthesised cartilage matrix increasing interface strength further [25]. This is the one place where a printed construct engages the stiffness of bone at all, and it does so mechanically, through anchorage, rather than by matching modulus.

Zonal Architecture and Matrix Maturation

Zonal printing attempts to reproduce the graded organisation of the osteochondral unit. Printability itself is now well characterised: gelatin-methacryloyl from 3% to 20% w/v combined with 0% to 1.5% gellan gum prints with cells between 15 and 37 degrees C, with yield stress the dominant determinant, and high gellan concentrations compromise cartilage matrix production before they prevent encapsulation altogether [26]. That trade-off — the same one seen in reinforcement — sets the ceiling on what a single-phase bioink can deliver.

Biphasic printing exposes a cost that monophasic work hides. In interwoven gel and calcium phosphate cement constructs, day-1 viability was 46.5 +/- 1.3% at the gel-cement interface and 53.8 +/- 2.5% at gel-gel interfaces, remaining near 50% to day 21, whereas monophasic alginate-methylcellulose scaffolds gave 65-75% early and 53% at day 21 under differentiation conditions; sGAG production reached approximately 10 mg/L per 1 x 10^6 cells and the cell number ratio at day 21 versus day 0 was approximately 1.0, indicating no net proliferation over three weeks [21]. Placing a mineral phase next to cells therefore costs roughly half of them, and the surviving population does not expand.

Phenotypic stability is the other unresolved axis. Tyrosinase-crosslinked silk fibroin-gelatin bioinks were used to interrogate hypertrophic regulation in printed hMSC constructs over 21 days, tracking COL10A1 and MMP13 alongside HIF1A and HDAC4 [27]. Gradient printing through a microfluidic head produced adjacent hyaline and calcified cartilage layers from hBM-MSCs and chondrocytes, with hypertrophic markers upregulated in the homogeneous calcified-cartilage equivalent but not in the gradient construct [28]. Neither line of work reports a modulus, so a phenotypically correct construct and a mechanically competent one have not yet been the same construct. Degradation completes the difficulty: no quantitative mass-loss or half-life data are available for any of these bioinks, and the only defensible statement is the qualitative one that hydrogel persisted at six months in a large-animal defect where cartilage did not form.

Large-Animal Evidence and the Translation Gap

The in vivo record is small, heterogeneous, and mechanically silent. In a sheep model, a standardised critical-sized full-thickness chondral defect was created on the weight-bearing surface of the lateral and medial condyles of both femurs in six animals, giving 24 defects across four arms — handheld in situ printing, bench-printed scaffold, microfracture and untreated control — and assessed at 8 weeks, with early cartilage regeneration in the handheld group and no operative complications [29]. In rabbits, a dual-factor-releasing gradient construct was compared across three groups with both knees used, scoring significantly higher on ICRS than non-gradient groups (p < 0.05, N = 6 per group) at euthanasia points of 8, 12 and 24 weeks; the compressive moduli were described as similar to native cartilage but the values sit in supplementary material [22].

Longer follow-up is less encouraging. A caprine osteochondral model at 6 months gave more hyaline-like repair with a biphasic fibre-reinforced construct than with commercial controls, but the quality of repair was variable from animal to animal [30]. An equine model at 6 months showed significant bone ingrowth into the printed PCL anchor and only limited production of cartilage-like tissue despite persistency of the hydrogel, with no difference in histological scores between zonal and non-zonal constructs [31]. Multilayered atelocollagen with supramolecular hyaluronic acid carrying human MSCs in a rabbit knee is reported favourably by its authors without supporting numbers [32]. Interspecies differences in cartilage mechanics are themselves substantial and site-dependent [33], so results from goat, horse, sheep, rabbit and rat models cannot be aggregated into a single efficacy claim. Reviewing barriers to translation, Francis and colleagues recorded that as of 2018 there were still no human clinical trials using 3D-printed engineered tissue for cartilage [34]; the persistence of the hydrogel and the absence of load-bearing stiffness in the equine result indicate why.

The mechanical gap is the organising fact of this field. Printed MSC-laden constructs measured in bulk compression span 5.2 kPa at day 0 to 42.0 kPa at day 56; native cartilage is quoted at 400-800 kPa by the same method and human subchondral bone at 13.34 +/- 1.69 GPa by nanoindentation. Fibre reinforcement closes the chondral part of that gap, reaching 405 +/- 68 kPa in an acellular composite, and interlocking architectures anchor a soft compartment to a stiff one, but neither achievement has yet been delivered together with the cell viability, phenotypic stability and matrix accumulation that a cell-laden construct requires. Progress will be legible only if every viability figure is published with its timepoint and modality and every modulus with its measurement method, so that a construct improving in compression is not credited with closing a nanoindentation gap it never touched.

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